Blood pump systems and methods

ABSTRACT

A blood pump system for persistently increasing the overall diameter and lumen diameter of peripheral veins and arteries by persistently increasing the speed of blood and the wall shear stress in a peripheral vein or artery for a period of time sufficient to result in a persistent increase in the overall diameter and lumen diameter of the vessel is provided. The blood pump system includes a blood pump, blood conduit(s), a control system with optional sensors, and a power source. The pump system is configured to connect to the vascular system in a patient and pump blood at a desired rate and pulsatility. The pumping of blood is monitored and adjusted, as necessary, to maintain the desired elevated blood speed, wall shear stress, and desired pulsatility in the target vessel to optimize the rate and extent of persistent increase in the overall diameter and lumen diameter of the target vessel.

CROSS REFERENCE TO RELATED APPLICATIONS

This application claims priority to U.S. Patent Application No.61/564,671 entitled “Blood Pump Systems and Methods,” filed on Nov. 29,2011, and claims priority to U.S. Patent Application No. 61/524,761,entitled “Blood Pump Systems and Methods,” filed on Aug. 17, 2011, whichis a continuation-in-part of U.S. patent application Ser. No.13/030,054, entitled “System and Method to Increase the Overall Diameterof Veins” filed on Feb. 17, 2011, which claims priority to U.S.Provisional Application No. 61/305,508 entitled “System and Method toIncrease the Overall Diameter of Veins” filed on Feb. 17, 2010, and isrelated to co-pending, co-filed PCT International Patent Application No.PCT/US12/50978, entitled “System and Method to Increase the OverallDiameter of Veins and Arteries,” filed on Aug. 15, 2012, and is relatedto co-pending U.S. Patent Application No. 61/524,759 entitled “Systemand Method to Increase the Overall Diameter of Veins and Arteries,”filed on Aug. 17, 2011, and U.S. Patent Application No. 61/561,859entitled “System and Method to Increase the Overall Diameter of Veinsand Arteries,” filed on Nov. 19, 2011, all of which are incorporated byreference in their entireties.

FIELD OF THE INVENTION

The present invention relates to a blood pump system that includes apump, conduits, a control unit, and a source of power, whereby thesystem may be used to persistently increase local blood flow in arteriesand veins of patients. Specifically, this invention may be useful forpersistently increasing the overall diameter and lumen diameter of veinsand arteries in patients needing a vascular access site forhemodialysis, a bypass graft, or other type of surgery or procedurewhere a larger vein or artery diameter is desired. This invention mayalso be useful for providing increased local blood flow to organs andtissues in need thereof, such as the lower extremities of patients withperipheral arterial disease (PAD).

BACKGROUND INFORMATION

There are over half a million chronic kidney disease (CKD) patients inthe United States, with over 100,000 new CKD patients each year. Thereis a four percent annual increase in projected prevalence population dueto such driving factors as, for example, high blood pressure, diabetes,and an aging population.

Hemodialysis is the treatment of choice for 92% of CKD patients, becausewithout hemodialysis or some other form of treatment those CKD patientswould die. A typical CKD patient undergoing hemodialysis treatment musthave his or her vascular system connected to a hemodialysis machine twoto three times per week. For hemodialysis, there are three commonvascular access site options. The preferred access site option is anarteriovenous fistula (AVF), which is a direct, surgically createdconnection between an artery and a vein, preferably in the wrist, oralternatively, in the forearm, upper arm, leg, or groin. Another accesssite option is an arteriovenous graft (AVG), which is a surgicallycreated connection between an artery and vein using an interposedsynthetic conduit. The final major access site option is a catheterinserted into a large vein in the neck, chest, leg, or other anatomiclocation.

Patients with an AVF have less morbidity, less mortality, and a lowercost of care compared with patients with an AVG or a catheter;therefore, an AVF in the wrist is the preferred form of vascular accessfor hemodialysis. Patients with an AVG or catheter have substantiallyhigher rates of infection and death than patients having an AVF, withcatheter patients having the worst outcomes. In addition, patientshaving an AVG or catheter have a higher average cost of care, withcatheter patients having the highest costs. If a patient is eligible foran AVF, the wrist or forearm is generally preferred over an AVF in theupper arm due to higher rates of hand ischemia and the generally shorterand deeper vein segments of the upper arm.

Unfortunately, about 85 percent of patients are ineligible for an AVF inthe wrist, mostly due to vein and artery diameters that are too small.Furthermore, about 60 percent of all AVFs created are not useablewithout additional surgical and interventional procedures due to anoccurrence commonly referred to as “maturation failure,” which iscorrelated with small vein and artery diameter. The availability ofveins and arteries with larger diameters is correlated with higher AVFeligibility and lower rates of maturation failure.

Currently, there are few options for permanently and persistentlyincreasing the diameter of a vein or artery. All current methods usemechanical methods of dilation, such as balloon angioplasty, that canlead to vein or artery injury. Since a patient needs to have peripheralveins and arteries of a certain size for a physician to create an AVF,it is desirable to have a method and system for persistently andpermanently increasing the size or diameter of peripheral veins orarteries.

Currently, small “heart pumps” exist. However, such pumps are costly andnot designed and dimensioned for use in an extremity. As such, there isa need in the art for systems, components, and methods of increasing thediameter of peripheral veins and arteries at a reasonable cost.Additionally, there is a need for a pump device that can increase thediameter of peripheral veins and arteries.

SUMMARY OF THE INVENTION

The present application relates to a blood pump system for use inincreasing the diameter of veins and arteries, preferably peripheralveins and arteries. The system will function to move blood in such a wayas to cause an increase in vein or artery diameters. This can beaccomplished by discharging (“pushing”) blood into a vein or artery orby removing (“pulling”) blood from a vein or artery. By either method,the system increases the flow of blood in a vessel, which ultimatelyleads to a persistent increase in vessel diameter. As such, the systemand, more particularly, the pump use mechanical means to activatebiological response pathways resulting in the enlargement or“remodeling” of veins or arteries. The system has a blood pump, conduitsto carry blood to and from the blood pump, a control system to monitorthe blood pump and modify the operation of the blood pump, and a powersource. As such, the system comprises a group of members that can be,for example, inserted into an artery at one end and a vein at the other,whereby, when activated, blood is pumped at a rate such that wall shearstress (WSS) on the endothelium of the vein, artery, or both is elevatedfor a period of time sufficient to causes a persistent enlargement inthe vein or artery. Any of a variety of pumps may be used so long as thepump can be controlled to produce the desired blood vessel diameterincrease.

Various types of blood pumps may be employed, including positivedisplacement and rotary pumps, with rotary type pumps being preferred.In one embodiment, a rotary blood pump system includes a pump having ahousing defining an inlet to receive blood and an outlet to dischargeblood. The pump housing is designed and dimensioned to house a rotatingimpeller suspended on bearings. The pump housing can have a firstbearing at the inlet portion of the housing and a second bearing at theoutlet portion of the housing. Blood enters and exits the rotatingimpeller, whereby the impeller increases the exit speed of the blood.This increased speed is recovered or translated as increased pressure asthe blood decelerates within the pump diffuser, which terminates in thepump outlet.

In other embodiments, various types of rotary blood pumps may be used.For example, an axial flow pump, a mixed flow pump, or preferably, acentrifugal blood pump may be used. In addition, a variety of pumpimpeller bearings may be used, including, but not limited to magneticbearings, hydrodynamic bearings, and, preferably pivot (contact) types.Similarly, various types of pump diffusers may be used, including butnot limited to a collector diffuser, or preferably a volute diffuser.

In one embodiment, a centrifugal blood pump with pivot bearings includesa pump housing defining a pump inlet having an inflow diffuser toreceive blood and direct blood onto an impeller, the pump housing havinga top bezel and top pivot bearing extending from a top of the housinginto the inlet, and a bottom bezel and bottom pivot bearing extendingfrom a bottom of the housing into the interior space of the housing. Thepump also includes the impeller suspended within the housing, theimpeller further having a bearing lumen to receive an impeller pivot.The impeller pivot has a first end to engage the inlet portion (top)pivot bearing and a second end to engage the outlet portion (bottom)pivot bearing. In one embodiment, the ends of the impeller pivot areconvex and at least one end of each pivot bearing is concave. In anotherembodiment, the ends of the impeller pivot are concave and the pivotbearings are convex. The impeller can include a variety of fin or bladeconstructions designed to contact and accelerate blood into the volute.For example, the impeller defines a plurality of blades on the topsurface of the impeller and extending radially from a center of theimpeller to an outer edge of the impeller. The blades accelerate bloodfrom the impeller's central inlet to its peripheral outlet. In anotheroption, the impeller does not include blades or fins, but does includemeans to move or propel blood. The impeller optionally includes at leastone washout lumen, cut-away, or bore extending parallel to a centralaxis of the impeller from a bottom surface through the impeller to a topsurface. The lumen is designed to prevent stagnation of blood under theimpeller and around the bottom pivot bearing.

The blood pump includes a motor, preferably electric, designed toactuate the impeller. In one embodiment, the blood pump includes a drivemotor having at least one magnet mechanically attached to the impellerand at least one armature mechanically attached to the housing. Thearmature induces an electromotive force on the at least one magnetattached to the impeller. The pump motor can be an axial-gap brushlessdirect current (DC) torque motor with sensorless back electromotiveforce (back emf) commutation. The motor employs a sintered alloy ofneodymium iron boron (NdFeB) for the magnets in the rotor and a 3-phaseplanar “racetrack” coil configuration in the stator. The motor has apancake aspect ratio, with a very small axial length in comparison toits diameter.

The blood pump system has one or more conduits including a first(inflow) conduit having two ends, a first end that is fluidly connectedto a location in the vascular system and receives blood from thatlocation, and a second end that is fluidly connected to the pump. Theinflow conduit delivers blood to the pump. The blood pump system has asecond (outflow) conduit having two ends, a first end that is fluidlyconnected to the pump and receives blood from the pump, and a second endthat is fluidly connected to a location in the vascular system. Theoutflow delivers blood to a location in the vascular system.

In various embodiments, the conduits of the blood pump system have anindividual length of between 2 cm and 110 cm and a combined lengthbetween 4 cm and 220 cm, and may be trimmed to a desired length by asurgeon or other physician, including during implantation of the pumpsystem. The conduits each have an inner diameter between 2 mm and 10 mm,and preferably between 4 mm and 6 mm. The conduits may be formed atleast in part from polyurethane (such as Pellethane® or Carbothane®),polyvinyl chloride, polyethylene, silicone elastomer,polytetrafluoroethylene (PTFE), expanded polytetrafluoroethylene(ePTFE), polyethylene terephthalate (PET, e.g. Dacron), and combinationsthereof. The conduits may further include an elastic reservoir.

All or portions of the conduits may be reinforced with a braided orspiral coiled shape memory material, such as nitinol, or otherself-expanding or radially expansive material. The conduits may havechamfered ends that fluidly connect to the vascular system. The ends canbe chamfered at an angle between 10 degrees and 80 degrees. One or moreof the conduits may have a number of holes or fenestrations in the wallsof the distal ends, when configured for placement within the lumen of ablood vessel or other intravascular location. The conduits may besecured to the pump using radially-compressive connectors.

In one embodiment, a blood pump system includes a blood pump and acontrol system to monitor the blood pump system and modify the operationof the blood pump to maintain an increased mean wall shear stress withinan artery or vein fluidly connected to the blood pump. The controlsystem is further configured to maintain mean wall shear stress within avein in the range of 0.76 to 23 Pa, or preferably in the range of 2.5 to10 Pa. In another embodiment, the control system monitors and maintainsan increased mean blood speed within an artery or vein fluidly connectedto the blood pump. In this embodiment, the control system is configuredto maintain mean blood speed within an artery or vein in the range of 10cm/s and 120 cm/s, or preferably in the range of 25 cm/s and 100 cm/s.In either embodiment, the blood pump system is configured to maintainincreased mean wall shear stress or increased mean blood speed for atleast 1 day, 7 days, 14 days, 28 days, 42 days, 56 days, 84 days, or 112days.

The blood pump system has a control system to achieve and maintain thedesired flow rate, which can optionally include a control device forreceiving information and controlling the operation of the pump of theblood pumping system. At a minimum, the control system can be manuallyactuated to adjust speed of the motor. Alternately, an automatic (i.e.“smart”) control system can be used. Optionally, the control systemincludes sensors that can be located in the pump, the conduits, or inthe vascular system of the patient. The control device can measure therotational speed of the motor based on the zero-crossings of theback-emf waveform. These zero crossings indicate magnetic pole reversalsof the rotor. The speed of the motor is controlled by pulse widthmodulation (PWM) of the input voltage, and torque is controlled by PWMof the input current. The control device also monitors other statevariables of the pump motor, such as current and voltage, from whichboth the flow rate through the blood pumping system and the wall shearstress in the peripheral blood vessel can be estimated and controlled.The control device preferably includes a memory, a processor forcontrolling the pump motor speed, analyzing the information coming fromthe motor drive electronics and optional sensors, and executinginstructions encoded on a computer-readable medium. The blood pumpsystem includes a cable for electrically connecting the control deviceto the pump and optional sensors. The blood pump system also includes apower source that, in various embodiments, may be integrated into thecontrol device. In various embodiments, the power source for the bloodpump system may be mobile (e.g. a rechargeable battery or fuel cell) orstationary (e.g. a power base unit connected to AC mains).

The control system may acquire information from various sources. Themotor drive electronics within the control device can measure at leastone of the motor speed, input power, or current required to operate thepump. In other embodiments, the control system includes sensors in theblood pump or conduits that measure at least one of a blood velocity, ablood flow rate, a resistance to blood flow in a peripheral bloodvessel, a blood pressure, a pulsatility index, and combinations thereof.In other embodiments, the control system includes sensors in thevascular system of the patient that measure at least one of a bloodvelocity, a blood flow rate, a blood pressure, a pulsatility index, avessel diameter, and combinations thereof.

In various embodiments, the control system may estimate and maintain adesired and elevated level of wall shear stress in a target vessel or adonating artery or vein, using the information from the control deviceand/or sensors, such as a motor speed, motor input power, pump flowrate, pump pressure head, pressure near the junction of the outflowconduit, and the target vessel, pressure drop across a blood vessel, andcombinations thereof. For the purpose of this application, “targetvessel”, “target blood vessel”, “target vein”, or “target artery” refersto a specific segment of an artery or a vein that is intended to achievea persistently increased overall diameter and lumen diameter when apump-conduit assembly is implanted, configured, and operated in such amanner as to result in the persistent increase in the overall diameterand lumen diameter.

Various control system methods may be used to automatically control theoperation of the blood pump system. In one embodiment, a method ofdetermining and controlling a wall shear stress in a blood vesselincludes the steps of measuring a blood viscosity, measuring a bloodflow rate in a blood pump system or the blood vessel, and measuring aradius of the blood vessel. The steps also include determining the wallshear stress in the blood vessel from the measured blood viscosity, themeasured flow rate, and the radius of the blood vessel, comparing thedetermined wall shear stress to a predetermined reference value, andadjusting a blood pump speed when the determined wall shear stress doesnot approximate the predetermined reference value. The steps arerepeated until the determined wall shear stress approximates thepredetermined reference value.

In another embodiment, a method of computing and controlling a wallshear stress in a blood vessel includes the steps of estimating a bloodviscosity, measuring a blood flow rate in a blood pump system or theblood vessel, and measuring a radius of the blood vessel. The steps alsoinclude determining the wall shear stress from the estimated bloodviscosity, the measured blood flow rate, and the radius of the bloodvessel, comparing the determined wall shear stress with a predeterminedreference value, and adjusting a blood pump speed when the determinedwall shear stress does not approximate the predetermined referencevalue. The steps are repeated until the determined wall shear stressapproximates the predetermined reference value.

In one embodiment, a method of estimating and controlling a wall shearstress in a blood vessel includes the steps of estimating a bloodviscosity, measuring at least one motor state variable of a blood pumpsystem selected from a voltage, a current, or a pump speed, andestimating a blood flow rate in the blood pump system. The steps alsoinclude measuring a pressure in the blood vessel, determining a vascularresistance of the blood vessel from the estimated blood flow rate andthe measured pressure in the blood vessel, estimating a radius of theblood vessel. The steps further include determining the wall shearstress from the estimated blood viscosity, the estimated blood flowrate, and the radius of the blood vessel, comparing the determined wallshear stress with a predetermined reference value, and adjusting thepump speed when the determined wall shear stress does not approximatethe predetermined reference value. The steps are repeated until thedetermined wall shear stress approximates the predetermined referencevalue.

In another embodiment, a method of estimating and controlling a wallshear stress in a blood vessel using a blood pump system includes thesteps of estimating a blood viscosity, measuring at least one motorstate variable of the blood pump system selected from a voltage, acurrent, or a pump speed, and estimating a blood flow rate and apressure head in the blood pump system. The steps also includecalculating a vascular resistance of the blood vessel from the estimatedblood flow rate and the estimated pressure head, estimating a radius ofthe blood vessel, and determining the wall shear stress from theestimated blood viscosity, the estimated blood flow rate, and the radiusof the blood vessel. The steps further include comparing the determinedwall shear stress with a predetermined reference value and adjusting thepump speed when the determined wall shear stress does not approximatethe predetermined reference value. The steps are repeated the determinedwall shear stress approximates the predetermined reference value.

In one embodiment, a method of estimating and controlling a wall shearstress in a blood vessel using a blood pump system includes the steps ofestimating at least one member selected from a group consisting of ablood viscosity, a blood flow rate, a pressure head in the blood pumpsystem, and a radius of the blood vessel, measuring at least one motorstate variable of the blood pump system selected from a group consistingof a voltage, a current, and a pump speed, and determining the wallshear stress in the blood vessel. The steps also include comparing thedetermined wall shear stress with a predetermined reference value andadjusting the pump speed when the determined wall shear stress does notapproximate the predetermined reference value. The steps are repeateduntil the determined wall shear stress approximates the predeterminedreference value.

In yet another embodiment, a sensorless method to avoid a collapse of ablood vessel fluidly connected to a blood pump system upon detecting animminence of the collapse at an inlet of the blood pump system includesthe steps of measuring a blood pump motor current and continuallydetermining a spectral analysis representation of the blood pump motorcurrent in a form of a Fourier series. The steps also include providinga detection indication when an amplitude of the second harmonic term ofthe Fourier series exceeds a reference value and decrementing a pumpspeed when the amplitude of the second harmonic term of the Fourierseries exceeds the reference value. The steps are repeated until theamplitude of the second harmonic term falls below the reference value.

In various other embodiments, the systems and methods disclosed hereinmay be encoded on computer-readable media that may be executed by a anyreference values or predetermined standards used by the systems andmethods may be stored in a database or other suitable storage medium.

BRIEF DESCRIPTION OF FIGURES

FIG. 1 is an isometric view of the pump.

FIG. 2 is an exploded isometric view of the pump showing its componentscontained in the body identified in FIG. 1.

FIGS. 3A and 3B are, respectively, partial and full cross sectionalelevations of the pump as taken along section line 3-3 in FIG. 1.

FIGS. 4A and 4B are, respectively, partial and full cross sectionalelevations of the pump as taken along section line 4-4 in FIG. 1.

FIGS. 5A-B are enlarged views of the pivot axis area of FIGS. 3B and 4B.

FIGS. 6A-B, respectively, are top and bottom isometric views of theimpeller pivot.

FIGS. 7A-B, respectively, are top and bottom isometric views of theimpeller pivot

FIGS. 8A-B are side elevation views of embodiments of the impellerpivot.

FIGS. 9A-B are, respectively, opposite end views of a representativebearing pin used on either end of the impeller pivot to support andallow rotation of the impeller pivot.

FIG. 10 is a view of an embodiment of the top bearing pin.

FIGS. 11A-B are side elevation views of embodiments of therepresentative bearing pin.

FIG. 12 is a longitudinal cross section of a representative bearing pinassembly.

FIG. 13 is a plan view of the inlet cap and impeller casing.

FIGS. 14-16 are, respectively, cross sectional elevations taken alongsection lines 14-14, 15-15, and 16-16 in FIG. 13.

FIG. 17 is an isometric partial cross section of the impeller chamberinlet orifice.

FIGS. 18A and 18B are, respectively, a plan view of the inlet capportion defining the inlet channel and an end elevation view of thesame.

FIGS. 19A and 19B are the same respective views as FIGS. 18A and 18B,except of another embodiment.

FIGS. 20A and 20B are the same respective views as FIGS. 18A and 18B,except of another embodiment.

FIGS. 21-23 are the same views as FIG. 18A, except of three otherembodiments.

FIGS. 24A and 24B are, respectively, plan and side elevation views ofanother embodiment of the inlet cap and inlet channel similar to thatdescribed in FIG. 21, except further including an arcuate wedgedportion.

FIG. 25 is an isometric view of the pump with the top impeller casingremoved to reveal the impeller occupying the impeller chamber.

FIG. 26 is a perspective view of a blood pump system according to oneembodiment.

FIGS. 27A-27D are perspective views of the connection between the pumpand conduits according to one embodiment.

FIGS. 28A and 28B are perspective views of the connection between thepump and conduits according to one embodiment.

FIGS. 29A and 29B are perspective views of the connection between thepump and conduits that include a side port according to one embodiment.

FIGS. 30A and 30B are perspective views of the connection between thepump and conduits that include a septum according to one embodiment.

FIG. 31 is a view of the distal portion of the outflow conduit accordingto one embodiment.

FIGS. 32A and 32B are views of the intravascular portion of an inflowconduit according to one embodiment.

FIG. 33 is a schematic view of the pump system according to oneembodiment.

FIG. 34 is a schematic view of the pump system according to anotherembodiment.

FIG. 35 is a schematic view of a control systems according to oneembodiment.

FIGS. 36A-36D are flowcharts of control system methods according tovarious embodiments.

FIGS. 36E is a plot of anastomosis pressures and blood flow rates for anin vitro model of the pump system according to one embodiment.

FIGS. 36F-36H are flowcharts of control system methods according tovarious embodiments.

FIG. 37 is a schematic view of the pump system as applied to acirculatory system of a patient according to one embodiment.

FIG. 38 is a schematic view of the pump system as applied to acirculatory system of a patient according to a second embodiment.

FIG. 39 is a schematic view of the system without a pump as applied to acirculatory system of a patient according to a third embodiment.

FIG. 40 is a schematic view of the pump system as applied to acirculatory system of a patient according to a fourth embodiment.

FIG. 41 is a longitudinal cross section of the junction between theproximal segment and distal segment.

FIG. 42 is a plan view of a medical kit.

FIG. 43 is a schematic diagram of a pump system controlled according tooutflow pressure.

DETAILED DESCRIPTION OF THE INVENTION

The systems and components of the present application relate to a bloodpump system. More specifically, in various embodiments, the presentapplication relates to a blood pump designed and dimensioned todischarge blood into a target vessel or withdraw blood from a targetvessel in such a way and for such a period of time that the diameter ofthe target vessel (vein or artery) is persistently increased. Even morespecifically, the present application relates to a rotary blood pumpsystem configured to persistently increase the mean and/or peak bloodvelocity and mean and/or peak wall shear stress in selected segments ofveins or arteries for a period of time sufficient to persistentlyincrease the overall diameter and the lumen diameter of selectedsegments of veins or arteries. The term “persistent increase” or“persistent dilation” when used to describe dilation or an increase inthe overall diameter and lumen diameter of an artery or vein, is usedherein to mean that even if the pump is turned off, an increase in theoverall diameter or lumen diameter of a vessel can still be demonstratedwhen compared to the overall diameter or lumen diameter of the vesselprior to the period of blood pumping. That is, the overall diameter orlumen diameter of the vessel has become larger independent of thepressure generated by the pump. The blood pump system may therefore beuseful to certain patients, including CKD patients in need of a vascularaccess site for hemodialysis. The blood pump system can include a rotaryblood pump, one or more blood-carrying conduits, a control system, and apower source. The blood pump system withdraws blood from one location inthe vascular system and discharges blood to another location in thevascular system. During operation, such a blood pump system maypersistently increase mean and/or peak blood velocity and mean and/orpeak WSS in a target blood vessel to a level and for a period of timesufficient to persistently increase the overall diameter and lumendiameter of the target blood vessel. The system functions inconfigurations where blood is withdrawn from the target blood vessel orin configurations where blood is discharged into the target bloodvessel. Further, the system can be used simultaneously to increase thesize of the donating and receiving vessels.

The optional blood-carrying conduits can include an inflow conduit tocarry blood from a location in the vascular system (such as a donatingvein, a donating artery, or the right atrium) to the blood pump and anoutflow conduit to carry blood from the blood pump to a location in thevascular system (such as an accepting peripheral vein or artery, or anaccepting location such as the right atrium). The blood pump system alsoincludes a control system. A preferred control system is designed tocollect information on the operating parameters and performance of theblood pump system, and changes in the vascular system, such as changesin the diameter of a donating artery, donating vein, accepting artery,or accepting vein of a patient. The blood pump system is primarilyconfigured to pump a sufficient amount of blood such that a desired meanand/or peak wall shear stress (WSS) is achieved within a blood vesselsegment (the “target blood vessel” or “target vessel”) and for asufficient period of time such that the permanent or persistent overalldiameter and lumen diameter of the blood vessel segment is increased.The mean WSS can be calculated using the measured, estimated, or assumedvessel diameter and the measured, estimated, or assumed average bloodflow rate through the blood pump system.

The diameter of blood vessels can be determined by measuring thediameter of the void within the center of the blood vessel. For thepurpose of this application, this measurement is referred to as “lumendiameter”. The diameter of blood vessels can be determined by measuringthe diameter in a manner that includes the void within the center of theblood vessel and the wall of the blood vessel. For the purpose of thisapplication, this measurement is referred to as “overall diameter”. Theinvention relates to simultaneously and persistently increasing theoverall diameter and lumen diameter of a peripheral vein by moving blood(preferably with low pulsatility) into the peripheral accepting vein,thereby increasing the speed of the blood in the peripheral acceptingvein and increasing the WSS on the endothelium of the peripheralaccepting vein. Systems and methods are described wherein the speed ofthe blood in a peripheral accepting vein and the WSS on the endotheliumof the peripheral accepting vein is increased by using a pump. Systemsand methods are also described that withdraw or “pull” blood such thatthe speed of the blood and the WSS is increased in the donating vessel,either an artery or a vein. Preferably, the pump actively dischargesblood into the peripheral accepting vein, wherein the pumped blood hasreduced pulsatility, such as when the pulse pressure is lower than bloodin a peripheral artery.

To begin a detailed discussion of the blood pump 25 of the system 10,reference is made to FIG. 1, which is an isometric view of the bloodpump 25. In one embodiment, the blood pump 25 is a miniaturizedcentrifugal pump having a magnetic drive wherein the impeller of thepump is rotationally driven by rotating magnetic fields. For example,the rotating magnetic fields may be generated by energizing a number ofelectromagnets in a particular sequence. In another example, therotating magnetic fields may be generated by rotating a number ofpermanent magnets or energized electromagnets. The pump can have adiameter approximately equal to that of a coin on the order of, forexample, a United States quarter, a United States half dollar, or alarger coin. As shown in FIG. 1, the blood pump 25 includes a body 105,an inlet 110, an outlet 115, and a power cable 120. The power cable 120connects the blood pump 25 to the control device 21 of a control system14 and power source. The power source can be part of the control device21 or separate. The power cable allows for communication between thecontrol device 21 and the motor of the blood pump 25. The cable can alsobe used to transfer power from a power source to the motor or pump. Moreparticularly, the power cable 120 connects the electrical components ofthe magnetic drive inside the body 105 to an electrical power source(e.g., a battery).

The inlet 110 is capable of being fluidly coupled to the inflow conduit20 via a coupling arrangement (e.g., a barbed-end, a flange, and alocking collar). The inlet 110 provides a fluid pathway into the intakeregion (i.e. center) of the pump impeller. The intake region of theimpeller can be of a variety of constructions so long as blood isreceived out of the outlet at a speed greater than the intake. Theoutlet 115 is capable of being fluidly coupled to the outflow conduit 30via a coupling arrangement similar to the inlet (e.g., a barbed-end, aflange, and a locking collar). The outlet 115 provides a fluid pathwayfrom the outlet region (i.e. periphery) of the pump impeller.

As illustrated in FIG. 2, which is an exploded isometric view of theblood pump 25 showing its components contained in the body 105identified in FIG. 1, the blood pump 25 includes an inlet cap 125, a topbearing pin 130, a top impeller casing 135, an impeller 140, an impellerpivot 145, a magnet assembly 150, a magnet enclosure 155, a bottombearing pin 160, a bottom impeller casing 165, an electrical coilassembly 170, and a coil assembly enclosure lid 175. The inlet cap 125and top impeller casing 135 each include approximately half of the inlet110.

As shown in FIGS. 3A and 3B, which are, respectively, partial and fullcross sectional elevations of the blood pump 25 as taken along sectionline 3-3 in FIG. 1, the components mentioned with respect to FIG. 2generally sandwich together to form the pump. For example, as can beunderstood from FIGS. 2-3A, the inlet cap 125 and top impeller casing135 respectively include a top horizontally extending inlet portion 110Aand a bottom horizontally extending inlet portion 110B. Typically, theinlet and outlet are opposed and located in different planes. When theinlet cap 125 and top impeller casing 135 are sandwiched together, theydefine an inlet fluid channel 180 leading through the inlet 110 to theimpeller inlet orifice 185. The inlet cap 125 and top impeller casing135 respectively define approximately a top half and a bottom half ofthe channel 180. A seal groove 190 is defined in the top impeller casing135 adjacent to the border of the channel 180 and is adapted to receivea resilient fluid seal member for creating a fluid tight seal betweenthe inlet cap 125 and top impeller casing 135.

FIGS. 4A and 4B are, respectively, partial and full cross sectionalelevations of the blood pump 25 as taken along section line 4-4 inFIG. 1. As can be understood from FIGS. 2, 4A, and 4B, the top impellercasing 135 and bottom impeller casing 165 respectively include a tophorizontally extending outlet portion 115A and a bottom horizontallyextending outlet portion 115B. When top impeller casing 135 and bottomimpeller casing 165 are sandwiched together, they define an outlet fluidchannel 200 (i.e. volute) leading from the impeller chamber 205 to theoutlet 115. The top impeller casing 135 and bottom impeller casing 165respectively define approximately a top half and a bottom half of thechannel 200. A seal groove 211 is defined in the bottom impeller casing165 adjacent to the border of the channel 200 and impeller chamber 205and is adapted to receive a resilient fluid seal member for creating afluid tight seal between the top impeller casing 135 and bottom impellercasing 165.

As indicated in FIGS. 2-4B, the magnets 150 are a plurality of magnetsin the form of a ring or disk. The magnets 150 are located in the volumeof the magnet enclosure 155 and the volume of the impeller 140. Themagnet enclosure is received in the impeller. The magnet enclosure 155and the impeller 140 respectively form the bottom and top portions ofthe volume in which the magnets 150 are located. The magnet enclosure,magnets, and impeller are coupled together in a fixed integral assemblythat rotates as a unit within the impeller chamber 205. Alternativeconstructions can be used that cause rotation of the impeller.

As illustrated in FIGS. 2-4B, the electrical coil assembly 170 is aplurality of electrical coils 210 arranged in a circular arrangement onthe lower impeller casing and optionally capped by a support disk 215.The electrical coil assembly 170 is fixed within the coil chamber 220defined in the bottom impeller casing 165 and capped by the coilenclosure lid 175. An internal floor structure 225 separates theimpeller chamber 205 from the coil chamber 220. The electrical cable 120(see FIG. 1) extends through passage 230 in the bottom impeller casing165 to the coil chamber 220 and the coils 210. Electrical power suppliedto the coils 210 via the electrical cable 120 generates rotatingmagnetic fields, which act on the magnets 150 to cause the magnets, andthe impeller 140 coupled to the magnets to rotate. The impeller rotationcauses the impeller blades 235 to act upon the fluid (e.g., blood)present in the impeller chamber, resulting in momentum being transferredto the fluid that is recovered as a pressure increase in the outletfluid channel 200. The fluid is thus drawn into the inlet 110 at lowpressure and discharged from the outlet 115 at a higher pressure.

As shown in FIGS. 3A-4B, the pivot axis for the impeller 140, magnets150, and enclosure 155 is the impeller pivot 145. As depicted in FIGS.5A-B, the impeller pivot 145 is pivotally supported (i.e. restrained inall degrees of freedom except rotation about a single axis) via a topbearing pin 130 and a bottom bearing pin 160. The top bearing pin 130 isreceived and fixed in a cylindrical recess 240 in the inlet cap 125,while the bottom bearing pin 160 is received and fixed in a cylindricalrecess 245 in the bottom impeller casing 165. The impeller pivot 145extends through and is fixed to a center cylindrical opening 250 in theimpeller 140.

In one embodiment of the impeller assembly, the impeller pivot 145, thetop bearing pin 130, and the bottom bearing pin 160 are formed from highpurity alumina, such as CoorsTek® AD-998. In another embodiment of theimpeller assembly, the impeller pivot 145, the top bearing pin 130, andthe bottom bearing pin 160 are formed from silicon carbide toughenedalumina, such as Greenleaf® WG-300. In both embodiments, the dimensionsof the impeller pivot 145, the top bearing pin 130, and the bottombearing pin 160 are designed to limit the contact stresses topermissible levels for high purity alumina or silicon carbide toughenedalumina, respectively, in view of peak thrust loads generated byhydrostatic forces and shock loads. In another embodiment of theimpeller assembly, the impeller pivot 145 is formed from silicon carbidetoughened alumina, such as Greenleaf® WG-300 or from high purityalumina, such as CoorsTek® AD-998, while the top bearing pin 130, thebottom bearing pin 160, or both are formed from ultrahigh molecularweight polyethylene. Additionally, the geometry of each component of theimpeller assembly has been selected to limit fatigue and wear in orderto satisfy the safety and durability requirements of the system 10.

As illustrated in FIGS. 6A-7B, the impeller pivot includes an upperhemispherical convex bearing surface 255 and a bottom hemisphericalconvex bearing surface 260. As indicated in FIGS. 6A, 6B, and 8A, oneembodiment of the impeller pivot has an overall length L1 ofapproximately 10.15 mm, plus or minus 0.05 mm, and a pivot diameter D1of approximately 2 mm, plus or minus approximately 0.01 mm. The upperbearing surface 255 has a radius R1 of approximately 0.61 mm, plus orminus 0.02 mm and extends a length L2 past an adjacent lip 265 byapproximately 0.55 mm, plus or minus 0.02 mm. The lower bearing surface260 has a radius R2 of approximately 0.31 mm, plus or minus 0.02 mm andextends a length L21 past an adjacent lip 265 by approximately 0.55 mm,plus or minus 0.02 mm. Similarly, an alternate embodiment of theimpeller pivot 145, as indicated in FIGS. 7A, 7B, and 8B, has an overalllength L1 of approximately 10.15 mm, plus or minus 0.05 mm, and a pivotdiameter D1 of approximately 2 mm, plus or minus approximately 0.01 mm.The upper bearing surface 255 has a radius R1 of approximately 0.31 mm,plus or minus 0.02 mm and extends a length L2 past an adjacent lip 265by approximately 0.55 mm, plus or minus 0.02 mm. The lower bearingsurface 260 has a radius R2 of approximately 0.31 mm, plus or minus 0.02mm and extends a length L21 past an adjacent lip 265 by approximately0.55 mm, plus or minus 0.02 mm. Other sizes and dimensions may be useddepending upon the size and performance requirements of the pump. Thesizes are such that the resultant pump can be used in a patient toincrease the diameter of a vessel.

As can be understood from FIGS. 5A and 5B, the upper bearings pin 130and bottom bearing pin 160 generally have the same configuration, butare oppositely oriented. As depicted in FIGS. 9A-B, the top bearing pin130 and the bottom bearing pin 160, have a tea cup or hemisphericalconcave bearing surface 270 on one end and a generally planar surface275 on the opposite end. Similarly, FIG. 10 depicts a particularembodiment of the top bearing pin 130, which has a tea cup orhemispherical concave bearing surface 270 on one end and a generallyplanar surface 275 on the opposite end. In this embodiment, thehemispherical concave bearing surface 270 of the top bearing pin 130 hasa larger radius than the concave bearing surface on the bottom bearingpin 160.

As illustrated in FIG. 11A, one embodiment of the bearing pin 130, 160has an overall length L3 of approximately 7.5 mm, plus or minus 0.1 mm,a minimum pivot diameter D2 of approximately 2 mm, plus or minus 0.01mm, and a radius of approximately 0.6 mm at the edge near the bearingsurface 270. Near the non-bearing end 275 of the bearing pin 130, 160, agroove 280 extends circumferentially around the pin to provide amechanical interlock for bonding the bearing pin in place within theblood pump 25. Similarly, an alternate embodiment of the bearing pins130, 160, as illustrated in FIG. 11B, has an overall length L3 ofapproximately 7.5 mm, plus or minus 0.1 mm, a minimum pivot diameter D2of approximately 3 mm, plus or minus 0.01 mm, and a radius ofapproximately 0.2 mm at the edge near the planar end 275. Near thenon-bearing end of the bearing pin 130, 160 there is a groove 280circumferentially extending around the pivot used to provide amechanical interlock for bonding the bearing pin in place. Other sizesand dimensions may be used depending upon the size of the pump, thematerials of the bearing pin, and the forces acting on the bearing pin.

As can be understood from FIGS. 3B, 4B, and 5A-11B, the convex upperbearing surface 255 of the impeller pivot 145 is rotationally receivedagainst the concave bearing surface 270 of the top bearing pin 130, andthe convex lower bearing surface 260 of the impeller pivot 145 isrotationally received against the concave bearing surface 270 of thebottom bearing pin 160. Thus, the convex bearing ends 255, 260 of theimpeller pivot 145 are pivotally supported by complementary concavebearing surfaces 270 of the top and bottom bearing pins 130 and 160,respectively. Accordingly, the impeller assembly may freely rotate inthe impeller chamber 205 on the impeller pivot 145, which is supportedend to end with the bearing pins 130 and 160, in a configurationcommonly known as a “double pin bearing.”

In yet another embodiment of the impeller assembly, the impellerassembly is a composite of the impeller shaft 145, top bearing pin 130,and bottom bearing pin 160. The composite design is beneficial withregard to the simplicity, tolerances, and cost of the machined bearingcomponents. All of these constructions are designed to allow the motorto function in a continuous state for around a day to 1-12 weeks orlonger, without breakdown.

As illustrated in FIG. 12, the impeller shaft 145 comprises an impellerpivot body 146 and two impeller pivot inserts 147. The impeller pivotbody 146 comprises a machinable metal, such as stainless steel, and theimpeller pivot inserts 147 comprise a high purity alumina, such asCoorsTek AD-998, or a silicon carbide toughened alumina, such asGreenleaf WG-300. The impeller pivot inserts 147 are affixed to theimpeller pivot body 146 by an adhesive and/or an interference fit.Optionally, the chamber 146A may be filled with an adhesive or otherpotting material that is resistant to compression. The aforementionedcomposite configuration and materials can be applied to embodiments ofboth the top bearing pin 130 and bottom bearing pin 160, where the pininserts 148 engage the impeller pivot inserts 147. Optionally, thechambers 148A for each bearing pin 130 and 160, may be filled with anadhesive or other potting material that is resistant to compression.

The inlet cap 125 and its inlet channel 180 may have a variety ofconfigurations, depending on the embodiment of the blood pump 25. Forexample, the inlet cap 125 depicted in FIG. 2 is shown as beinggenerally coextensive with the top impeller casing 135. In otherembodiments, the inlet cap 125 may be substantially smaller than, andnot coextensive with, the top impeller casing 135, as depicted in FIGS.13-15, which are views of the inlet cap and impeller casing.

As shown in FIGS. 14-16, which are, respectively, cross sectionalelevations taken along section lines 14-14, 15-15, and 16-16 in FIG. 13,the inlet 110 is a two part construction having portions 110A and 110Bthat each form approximately half of the inlet 110 and are respectivelypart of the inlet cap 125 and top impeller casing 135. Each portion 110Aand 110B has defined therein approximately half of the inlet channel180. As illustrated in FIG. 14, the inlet channel 180 initially has acircular diameter D5 of approximately 4 mm. As indicated in FIG. 15, theinlet channel 180 transitions from a circular cross section to agenerally rectangular cross section having a width W5 of approximately8.4 mm and a height H5 of approximately 1.5 mm. Again, as dimensionschange so will the listed measurements.

As depicted in FIG. 16, the inlet channel 180 surrounds the impellerchamber inlet orifice 185, which extends around the top bearing 145received in, and affixed to, the inlet cap 125. As shown in FIG. 17,which is an isometric partial cross section of the impeller chamberinlet orifice 185, the impeller chamber inlet orifice 185 leads to theimpeller chamber 205 near the intake region 300 of the impeller 140. Theupper bearing end of the impeller pivot 145 extends up through theorifice 185 to pivotally interface with the top bearing pin 130supported in the inlet cap 125. Impeller blades 235 extend radiallyoutward from the intake region 300 of the impeller 140.

As depicted in FIGS. 18A and 18B, which are, respectively, a plan viewof the inlet cap portion 110A defining the inlet channel 180 and an endelevation view of the same, in one embodiment, the inlet channel 180 maybe said to have an elliptic configuration. Specifically, a cylindricalchannel portion 180A transitions in portion 180C into an ellipticalchannel portion 180B. A cylindrical island portion or bezel 305supporting the top bearing pin 130 is generally centered in theelliptical channel portion 180B and includes a cylindrical hole 240 thatreceives the top bearing pin 130 similar to as illustrated in FIG. 17.In one embodiment, the cylindrical channel portion 180A has a diameterD6 of approximately 4 mm. The elliptical channel portion 180B has awidth W6 of approximately 12.4 mm. The distal distance W7 between thewall of the bezel 305 and the distal end of the wall defining theelliptical channel portion 180B is approximately 1.5 mm. In otherembodiments, the cylindrical channel portion 180A has a diameter D6 ofapproximately 5 mm or 6 mm.

As depicted in FIGS. 19A and 19B, which are the same respective views asFIGS. 18A and 18B, except of another embodiment, the inlet channel 180may be said to have a circular configuration. Specifically, acylindrical channel portion 180A transitions in portion 180C into acircular channel portion 180B. A cylindrical island portion or bezel 305supporting the top bearing pin 130 is generally centered in the circularchannel portion 180B and includes a cylindrical hole 240 that receivesthe top bearing pin 130 similar to as illustrated in FIG. 17. In oneembodiment, the cylindrical channel portion 180A has a diameter D9 ofapproximately 3.5 mm to 4.5 mm, preferably 4 mm. The circular channelportion 180B has a width W9 of approximately 11.5 mm to 13 mm,preferably 12.4 mm. The distal distance W10 between the wall of thebezel 305 and the distal end of the wall defining the circular channelportion 180B is approximately 3.5 mm to 4.5 mm, preferably 4.2 mm. Inother embodiments, the cylindrical channel portion 180A has a diameterD6 of approximately 5 mm or 6 mm.

As depicted in FIGS. 20A and 20B, which are the same respective views asFIGS. 18A and 18B, except of another embodiment, the inlet channel 180may be said to have a complex arcuate configuration. Specifically, acylindrical channel portion 180A transitions in portion 180C into acomplex arcuate channel portion 180B. A cylindrical island portion orbezel 305 supporting the top bearing pin 130 is generally centered inthe complex arcuate channel portion 180B and includes a cylindrical hole240 that receives the top bearing pin 130 similar to as illustrated inFIG. 17. In one embodiment, the cylindrical channel portion 180A has adiameter D12 of approximately 4 mm. The complex arcuate channel portion180B has a width W13 of approximately 8.4 mm. The distal distance W14between the wall of the bezel 305 and the distal end dome 307 of thewall defining the complex arcuate channel portion 180B is approximately1.75 mm. The distal distance W15 between the wall of the bezel 305 andthe distal end cleft 310 of the wall defining the complex arcuatechannel portion 180B is approximately 0.5 mm to 1.5 mm, preferably 1 mm.In other embodiments, the cylindrical channel portion 180A has adiameter D6 of approximately 5 mm or 6 mm.

As depicted in FIGS. 21-23, which are the same views as FIG. 18A, exceptof three other embodiments, the inlet channel 180 may be said to have atear drop configuration. Specifically, a cylindrical channel portion180A transitions into a tear drop channel portion 180B. A cylindricalisland portion or bezel 305 supporting the top bearing pin 130 isgenerally centered in the tear drop channel portion 180B and includes acylindrical hole 240 that receives the top bearing pin 130 similar to asillustrated in FIG. 17. In one embodiment, the cylindrical channelportion 180A has a diameter D15 of approximately 4 mm. The tear dropchannel portion 180B has a width W20 of approximately 8 mm. The bezel305 has a diameter D16 of 4 mm. A transition region 180C of the channel180 between the tear drop portion 180B and the cylindrical portion 180Ahas walls that diverge from each other at an angle AN1 of approximately8 degrees. In other embodiments, the cylindrical channel portion 180Ahas a diameter D6 of approximately 5 mm or 6 mm.

For the embodiment of FIG. 21, the distal distance W21 between the wallof the bezel 305 and the distal end of the wall defining the tear dropchannel portion 180B is approximately 2 mm. For the embodiment of FIG.22, the distal distance W21 between the wall of the bezel 305 and thedistal end of the wall defining the tear drop channel portion 180B isapproximately 1 mm. For the embodiment of FIG. 23, the distal distanceW21 between the wall of the bezel 305 and the distal end of the walldefining the tear drop channel portion 180B is approximately 0 mmbecause the bezel intersects the distal end of the wall defining thetear drop channel portion.

As illustrated in FIGS. 24A and 24B, which are, respectively, plan andside elevation views of another embodiment of the inlet cap 110 andinlet channel 180 similar to that described in FIG. 21, an arcuatewedged portion 320 may extend between the distal wall of the tear dropchannel portion 180B to the distal side of the bezel 305. In such anembodiment, the cylindrical island portion or bezel 305 is generallycentered in the tear drop channel portion 180B and includes acylindrical hole 240 that receives the top bearing pin 130 similarly toas illustrated in FIG. 17. In one embodiment, the dimensionalconfiguration of the embodiment depicted in FIGS. 24A and 24B issubstantially the same as discussed with respect to FIG. 21, thesignificant difference being the presence of the arcuate wedge portion320. As can be understood from FIGS. 24A and 24B, the wedge portion 320has walls that are arcuate to smoothly curve from the roof and adjacentwall of the tear drop channel portion 180B to the vertical extension ofthe bezel 305. Such a wedged portion 320 may be seen to exist in theembodiment depicted in FIGS. 3A, 3B, and 17 and may reduce areas ofinlet channel flow stagnation and facilitate tangential inflow of fluidthrough the impeller chamber inlet orifice 185.

As shown in FIG. 25, which is an isometric view of the blood pump 25with the top impeller casing removed to reveal the impeller 140occupying the impeller chamber 205, the outlet fluid channel 200 exitsthe impeller chamber substantially tangential to the outercircumferential edge of the impeller. As indicated in FIGS. 3B, 4B, 17,and 25, a plurality of bores 350 (i.e. washout holes) arecircumferentially distributed about the impeller pivot center hole 250,and the bores 350 are generally parallel to the center hole 250 andextend though the full thickness of the impeller to daylight on both topand bottom boundaries of the impeller. The bottom openings of the bores350 are located near the bottom bearing interface between the bottombearing 165 and the impeller pivot bottom bearing surface 260 (see FIG.8). As a result, a fluid can be flowed through the bores 350 to cleansethe bottom bearing interface. For example, a fluid can be flowed throughthe impeller chamber inlet hole 185, radially-outward along the impellerblades 235, through the gap under the impeller, and then back to theregion of the impeller chamber inlet hole 185. This flow of blood servesto cleanse the underside of the impeller, the bottom bearing interface,the upper bearing interface, and the region behind the bezel 305.

As can be understood from FIGS. 3B, 5, 17, and 25, in one embodiment,the impeller 140 is rotationally supported in the impeller chamber 205on a shaft 145 extending through a center of the impeller. The shaft hasan upper bearing end and a bottom bearing end, each end rotatablyoperably coupled to the pump housing. The impeller has a top face, abottom face, and multiple bores 350 extending through the impeller fromthe top face to the bottom face. The multiple bores are generally evenlydistributed radially about center of the impeller. Further, the multiplebores extend through the impeller generally parallel to each other andthe shaft. The inlet channel 180 leads to an inlet orifice 185 of theimpeller chamber. The inlet channel opens into the impeller chambergenerally perpendicular to the inlet channel. The inlet orifice extendsalong at least a portion of an outer circumferential surface of theshaft near the upper bearing end. The inlet orifice and the holes openin directions that are generally parallel to each other. Duringoperation of the pump, at least a portion of the blood pumped throughthe impeller chamber circulates along the top and bottom faces of theimpeller via the bores. Thus, the bores of the impeller eliminate flowdead ends around the impeller by generally keeping blood flowing alongall blood contacting surfaces of the impeller. Accordingly, the boreshelp to prevent blood accumulation in the vicinity of the shaft/impellerintersection and along the sides and bottom face of the impeller.

The body and impeller of the blood pump 25, including blood-contactingsurfaces, are made from a variety of rigid biocompatible materials. Oneoption includes plastics, more preferably injection moldable plasticssuch as PEEK. In various embodiments, the blood-contacting surfaces ofthe blood pump 25 may comprise Ti₆Al₄V, Ti₆Al₇Nb, or other commerciallypure titanium alloys. In one embodiment, the surfaces of the pumpcomponents to be exposed to the patient's blood may have antithromboticcoatings. For example, the luminal surfaces may be coated with Astute®,a heparin based antithrombotic coating by BioInteractions Ltd., orApplause™, a heparin coating by SurModics, Inc.

In other embodiments, the surfaces of the blood pump system componentsin contact with the patient's tissue may have antimicrobial coatings.For example, the external surfaces of the synthetic conduits 16 and 18or the external surfaces of the pump or the power cord 120 (which isalso know as a “lead”) may be coated with Avert®, a surface-activeantimicrobial coating by BioInteractions Ltd.

In various embodiments, the blood pump 25 may be implanted within apatient. Conversely, in other embodiments, the blood pump 25 may remainexternal to the patient. For example, when located externally to thepatient, the blood pump 25 may be secured to the patient using tape,sutures, or other suitable means to affix the pump to the patient. Thesystem 10 may be powered by wearable electronics having rechargeablebatteries 28, as shown in FIG. 34.

The pump for the pump system 10 disclosed herein may be a rotary pump,including, for example, a centrifugal flow pump, an axial flow pump, aradial flow pump, or a mixed flow pump. As shown in FIGS. 1-15, in oneembodiment, the pump is a centrifugal pump. Without recognizing specificlimitations, the blood pump 25 can be configured to routinely pump about0.05 to 1.0 L/min, 0.2 to 1.5 L, or 0.5 to 3.0 L/min, for example.

While the pump configuration discussed above with respect to FIGS. 1-25is advantageous, other pump configurations may be employed with the pumpsystems and methods disclosed herein. Accordingly, the systems andmethods disclosed herein should not be limited to the pump configurationdiscussed above with respect to FIGS. 1-25, but should include all typesof pumps applicable for the systems and methods disclosed herein.

A preferred embodiment of the pump system 10 disclosed herein withrespect to FIGS. 1-25 satisfies several unique needs that cannot besatisfied by any blood pump systems known in the art. Specifically, theArteriovenous Fistula Eligibility (“AFE”) pump system (“AFE System”) maybe configured for up to 12 weeks of intended use. Further, the AFE pumpsystem may be configured as a centrifugal rotary blood pump system forlow flow rate (e.g., 50 to 1500 mL/min) and medium pressure range (e.g.,25 to 350 mmHg). A control scheme used with the AFE pump system may beoptimized to maintain a steady and elevated mean WSS of 0.76-23 Pa intarget veins that are directly fluidly connected to the blood pump or aconduit of the blood pump system, or target veins that are fluidlyconnected to a vein that is directly fluidly connected to the blood pumpor a conduit of the blood pump system. The AFE pump system is configuredto operate for a period of time such that the overall diameter and lumendiameter of the target vein will persistently increase by 25%, 50%, or100% or more, utilizing sensing of operating parameters and periodicspeed adjustment.

For certain embodiments, the inflow conduit may be placed bypercutaneous approach, with a portion of the inflow conduit residing inan intravascular location, and the outflow conduit may be placed bysurgical approach adaptable to initial vein diameters of between 1-6 mm.In this setting, elevated mean WSS in the target blood vessel resultsfrom discharging blood into the target blood vessel.

For other embodiments, the outflow conduit may be placed by percutaneousapproach, with a portion of the outflow conduit residing in anintravascular location, and the inflow conduit may be placed by surgicalapproach adaptable to initial vein or artery diameters of between 1-6mm. In this setting, elevated mean WSS in the target blood vesselresults from removing blood from the target blood vessel. In certainsettings, WSS can be elevated in both a blood vessel where blood isremoved and a blood vessel where blood is discharged, making both bloodvessels target blood vessels. The pump system 10 achieves both ease ofinsertion/removal and resistance to infection. The pump system 10 is amobile system with a pump that is adaptable for either implanted orextracorporeal placement. In various embodiments, the pump system 10 ispowered by wearable electronics with rechargeable batteries.

The pump system 10 includes an inflow conduit 20 and an outflow conduit30, as shown in FIG. 26. The inflow conduit 20 is placed in fluidcommunication with one location in the vascular system, draws blood fromthis location, and carries it to the blood pump 25. In certainembodiments, the inflow conduit 20 is configured for placement of atleast a portion of the inflow conduit within the lumen of the vascularsystem. In other embodiments, the inflow conduit 20 is joined to a bloodvessel by a surgical anastomosis. The outflow conduit 30 is configuredfor making a fluid communication with another location in the vascularsystem and directs blood from the blood pump 25 to the other location inthe vascular system. In certain embodiments, the outflow conduit 20 isconfigured for placement of at least a portion of the outflow conduitwithin the lumen of the vascular system. In other embodiments, theoutflow conduit 30 is joined to a blood vessel by a surgicalanastomosis.

The conduits 20 and 30 may each have a length that ranges between 2 cmand 110 cm and a total combined length of 4 cm to 220 cm. The length ofthe each conduit 20 and 30 may be trimmed to a desired length asdetermined by the location of the blood pump 25 and the location of theconnections between the conduits and the vascular system. The conduits20 and 30 also have thin but compression-resistant and kink-resistantwalls that have a thickness of between 0.5 mm and 4 mm and innerdiameters that are between 2 mm and 10 mm. Preferably, the innerdiameters for the conduits are 4 to 6 mm.

The inflow and outflow conduits 20 and 30 may be connected to the bloodpump 25 using any suitable connector that is durable, resists leaks, andis not susceptible to unintentional disengagement. Typically, theleading edge of the connector is thin, in order to minimize the stepchange in fluid path diameter between the inner diameter of the conduits20 and 30 and the inner diameter of the connector. Preferably, the stepchange in fluid path diameter should be less than 0.5 mm. In oneembodiment, as shown FIGS. 27A-27D, the conduits 20 and 30 are connectedto the blood pump 25 using barb fittings 400A and 400B and radiallycompressive retainers (i.e. locking collars) 402A and 402B. By way ofexample, and not limitation, the radially compressive retainers 402A and402B, may be BarbLock® retainers manufactured by Saint-GobainPerformance Plastics, a division of Saint-Gobain S.A. headquartered inCourbevoie, France. In another embodiment, the conduits 20 and 30 areconnected to the blood pump 25 using Pure-Fit® sterile connectors, alsomanufactured by Saint-Gobain Performance Plastics.

The radial compressive retainers 402A and 402B are placed over theproximal ends 404 and 406 of the inflow and outflow conduits 20 and 30,respectively. The conduits 20 and 30 are then placed over the barbfitting 400A and 400B to form a fluid connection between the conduitsand the blood pump 25. Collets 408A and 408B of the radial compressiveretainers 402A and 402B are placed along the conduits 20 and 30 toencircle the conduits and the barb-fittings 400A and 400B. Outer sleeves410A and 410B of the radial compressive retainers 402A and 402B are thenmoved along a longitudinal axis of the retainers to compressively engagethe respective collets 408A and 408B, conduits 20 and 30, and the barbfittings 400A and 400B. In one embodiment, the outer sleeves 410A and410B are moved by a compressive tool configured to engage the outersleeves and a support shelf 412A and 412B of the barb fittings 400A and400B, respectively. The compressive tool may also be configured toremove the radial compressive retainers 402A and 402B.

In other embodiments, alternative connectors may be used. Preferably,the alternative connectors are durable, resist leaks, and resistunintentional dislodgment. For example, as shown in FIG. 28A-B, theconduits 20 and 30 engage barb fittings, similar to barb fittings 400Aand 400B, to form a fluid connection between the conduits and the bloodpump 25. The conduits 20 and 30 are secured to the barb fittings usingcircular clips 414A and 414B that apply radial compressive force to theportion of the conduits on the barb fittings by way of a ratchetingmechanism 416A-416B of the clips. The circular clips 414A and 414Bprovide a leak-resistant and durable connection that may be removed witha removal tool (not shown) which releases the ratcheting mechanisms416A-416B of the clips.

In another embodiment, the inflow conduit 20 and the outflow conduit 30contain side ports that provide controlled access to the fluid path.Side ports may be used periodically to introduce contrast into the fluidpath to enable visualization by fluoroscopy, to obtain blood samples, toinfuse medications, or for other clinically useful purposes. Any sideport design that allows periodic access to the fluid path and does notleak or alter the fluid flow path when not accessed is suitable. By wayof example, and not limitation, the side port may be a “T” port fittingthat includes a check valve that opens when a syringe is inserted andcloses when the syringe is removed. As shown in FIGS. 29A-B, a “T” portassembly 418 with auxiliary tubing 420 is in fluid communication withthe pump outlet 115 and outflow conduit 30.

In another embodiment, a side port for the inflow conduit 20, theoutflow conduit 30, or both utilizes a septum access port 422 having aseptum 424, as shown in FIGS. 30A-B, through which a suitable hypodermicneedle can be inserted for access and then removed, after which theseptum closes, preventing fluid loss from the conduit. Suitablematerials for the septum 424 include, but are not limited to silicone,polyurethane, and other elastomeric polymers. The segment of the inflowand/or outflow conduit 20 or 30, respectively, which includes the septum424, is of a suitable thickness to close a hypodermic puncture hole whenthe needle is removed. As shown in FIGS. 30A-B, the septum access port422 is shown in which the septum 424 makes up a portion of the outflowconduit 30. By way of example, and not limitation, the septum accessport 422 may extend about one centimeter over the length of the outflowconduit 30. The septum 424 may be attached to the outflow conduit 30 byany suitable means including, but not limited to, adhesive attachment,thermal bonding, and thermal bonding between inner and outer layers ofthe conduit tubing.

In various embodiments, the conduits 20 and 30 may be comprised ofmaterials commonly used to make hemodialysis catheters such aspolyurethane, polyvinyl chloride, polyethylene, silicone, andpolytetrafluoroethylene (PTFE), and including Pellethane® orCarbothane®. In other embodiments, the conduits may be comprised ofmaterials commonly used to make hemodialysis grafts or syntheticperipheral bypass grafts such as expanded polytetrafluoroethylene(ePTFE) or Dacron. In further embodiments, conduits may be comprised ofcombinations of polyurethane, polyvinyl chloride, polyethylene,silicone, PTFE, Pellethane®, Carbothane®, Carbothane® PC-3575, ePTFE, orDacron.

For example, the entire length of the inflow conduit 20 may be composedof polyurethane. In another embodiment, shown in FIG. 31, a segment 500of the outflow conduit 30 configured to make a fluid communication withthe blood pump 25 is composed of polyurethane while a segment 502 of theoutflow conduit configured to make a fluid communication with thevascular system is composed of ePTFE.

By way of example and not limitation, and as shown in FIG. 41, which isa longitudinal cross section of the junction between the proximalsegment 500 and distal segment 502, the proximal segment 500 of theoutflow conduit 30 is joined to the distal segment 502 of the outflowconduit during the manufacturing process by placing one or more layers502A of ePTFE from the distal segment between layers 500A ofpolyurethane from the proximal segment. The overlapping layers ofpolyurethane and ePTFE are then heat laminated to bond the proximalsegment 500 and the distal segments 502 together.

In another example, one or more holes are made within the overlappedsections of the ePTFE of segment 502 prior to heat laminating theconduit. When the outflow conduit 30 is heated to a temperature that issufficient to melt the polyurethane without melting the ePTFE (e.g. 200°F. to 500° F.), the molten polyurethane fills in and then cools withinthe holes created in the ePTFE segment 502. The inner and outerpolyurethane layers of the segment 500 are joined with in the holes tomechanically join the two segments 500 and 502 together as well asmechanically join the inner and outer layers of polyurethane in theoverlapped segment.

The embodiment of the outflow conduit 30 manufactured to have the ePTFElayer 502A sandwiched between the polyurethane layers 500A isadvantageous in that the ePTFE layer 502A can be readily sutured toblood vessels using standard techniques. This is also the case for aninflow conduit 20 manufactured as discussed above with respect to FIG.41.

As illustrated in FIG. 42, which is a plan view of a medical kit 1000,the blood pump 25, inflow conduit 20, outflow conduit 30, control device21, and power cord 120 can be provided in a sterile package 1005 withinstructions 1010 on how to assemble and implant the pump system in apatient. The medical kit 1000 may also include the barb fittings 400Aand 400B and the radially compressive retainers 402A and 402B. In oneembodiment, one or both conduits 20, 30 are manufactured as describedabove with respect to FIG. 41 and enclosed within the sterile package1005 along with the blood pump 25. The medical kit 1000, at a minimum,includes a system for discharging or removing blood and instructions forimplementation and usage.

In one embodiment, the operation of the blood pump 25 is controlled viathe control unit 21 of a pump control system 14 by reading the outflowpressure and adjusting the pump speed accordingly. For example, asdepicted in FIG. 43, which is a schematic diagram of a pump system 10controlled according to outflow pressure, an outflow pressure sensor1050 may be operably coupled to the outlet 115 of the blood pump 25 orfurther downstream, such as, for example, somewhere along the length ofthe outflow conduit 30. The processor 24 may compare the pressurereading from the outflow pressure sensor 1050 to a range of targetoutflow pressures stored in the memory 27. The processor will thenadjust the speed of the pump drive 170 accordingly to cause the pressurereading from the outflow pressure sensor 1050 to be within the range oftarget outflow pressures stored in the memory.

In one embodiment, the control system 14 also includes an inflowpressure sensor 1060 that may be operably coupled to the inlet 110 ofthe blood pump 25 or further upstream, such as, for example, somewherealong the length of the inflow conduit 20. The processor 24 may readboth the pressure reading from the outflow pressure sensor 1050 and thepressure reading from the inflow pressure sensor 1060 and calculate apressure difference. This pressure difference may then be compared to arange of target pressure differences stored in the memory 1055. Theprocessor will then adjust the speed of the pump drive 170 to cause thecalculated pressure difference to be within the range of target pressuredifferences stored in the memory.

In other embodiments, the inflow and outflow conduits 20 and 30 can beany material or combination of materials so long as the conduits 20 and30 exhibit desirable characteristics, such as flexibility, sterility,resistance to kinking and compression, and can be connected to a bloodvessel via an anastomosis or inserted into the lumen of a blood vessel,as needed. In addition, the conduits 20 and 30 preferably exhibit thecharacteristics needed for subcutaneous tunneling as desired, such ascomprising lubricious external surface coatings such as Harmony™advanced lubricity coatings.

As another example, the inflow and outflow conduits 20 and 30 may havean exterior layer composed of a different material than the interiorlayer. All or a portion of the external layers of the inflow and outflowconduits 20 and 30 may also be coated with a lubricating agent, such assilicon or a hydrophilic coating to aid in subcutaneous tunneling andremoval from the body, and to mitigate possible allergic reactions tolatex. In certain embodiments, at least a portion of the surface of theexterior layer of the inflow and outflow conduits 20 and 30 may have anantimicrobial coating. In other embodiments, at least a portion of thesurface of the blood pump 25 or the power cord 120 may have anantimicrobial coating. For example, Avert™, a surface activeantimicrobial coating may be used. In certain embodiments, a portion ofthe surface of the exterior layer of an inflow and outflow conduit mayinclude a material to resist infection and encourage tissueincorporation, such as Dacron velour, polyester velour, or silicone. Onesuch material is the VitaCuff® antimicrobial cuff by Vitaphore Corp. TheVitaCuff cuff is comprised of two concentric layers of material. Theinternal layer is constructed of medical grade silicone. The external,tissue-interfacing layer comprises a collagen matrix with anantimicrobial activity that is attributable to silver ions bound to thecollagen. In certain embodiments, this material absorbs physiologicalfluids, quickly expands, and helps provide a physical barrier at theexit site. Tissue in-growth occurs, further securing the conduit inplace, and reducing conduit movement to reduce the incidence of exitsite infection.

In certain embodiments, at least a portion of the blood-contactingluminal surfaces of the inflow and outflow conduits 20 and 30 may becoated with an antithrombotic agent or material. Similarly, at least aportion of the blood-contacting surfaces of the blood pump 25 may becoated with an antithrombotic agent or material. For example, thesurfaces may be coated with the Applause® coating from SurModics, Inc.,or the Astute® coating from BioInteractions Ltd., which are bothhydrophilic copolymer coatings containing heparin.

In certain embodiments, at least a portion of the inflow conduit 20 andoutflow conduit 30 are preferentially reinforced to resist kinking andcompression. For example, the conduits 20 and 30 may be reinforced withnitinol or another shape memory alloy or self-expanding or radiallyexpansive material. Preferably, a layer of braided nitinol is wrappedaround at least a portion of each of the conduits 20 and 30 orincorporated into the walls of conduits. In one embodiment, the inflowconduit 20 is reinforced by braided nitinol incorporated into the wallsof the conduit. In another embodiment, the inflow conduit may bereinforced by braided stainless steel that is incorporated into the wallof the conduits 20 and 30. Alternately, a coil of nitinol or PTFE may bewrapped around portions of the conduits 20 and 30 or incorporatedtherein. For example, as shown in FIG. 31, the distal segment 502 of theoutflow conduit 30 has a PTFE coil 504 incorporated around the ePTFEconduit forming the wall 514 of the conduit. In other embodiments, acoil of nitinol may be wrapped around portions of the conduits 20 and 30or incorporated therein.

The braid density of the braided nitinol incorporated into both theinflow and the outflow conduits 20 and 30, commonly measured in pixelsper inch (“PPI”), is typically between about 10 and 200, and preferablybetween about 20 and about 60. In various embodiments, the braid densitymay vary along the lengths of the inflow and the outflow conduits 20 and30. For example, the braid density may be greater in portions of theconduits 20 and 30 adjacent to the blood pump 25, in order to maintaingreater stiffness of the conduits and minimize the risk of externalconduit compression or conduit collapse during suction, while allowingfor more flexibility in different segments of the conduits.

In one embodiment, as shown in FIGS. 32A-32B, the intravascular portion506 of the inflow conduit 20 is fenestrated by means of multiple sideholes 508. These side holes enhance blood inflow and reduce the risk ofsuction of the vein or right atrium wall by the end hole in the event ofpartial obstruction of the conduit tip. Preferably, the side holes 508are circular and range in diameter from 0.5 mm to 1.5 mm. In otherembodiments, however, the side holes 508 may be elliptical or any othershape and size suitable for the intravascular aspiration of blood.

As shown in FIGS. 31 and 32A-32B, the distal end 506 of the inflowconduit 20 and the distal end 510 of the outflow conduit 30 may be cutand chamfered at an angle between 10° and 80°. In certain embodiments,the chamfer reduces the risk of suction of the vein or right atrium wallby the end hole in the event of partial obstruction of the tip of theconduit during aspiration of blood. In other embodiments, the chamferincreases the area of the conduit as it joins the vascular system in ananastomotic connection. Preferably, but without limitation, the distalends 506 and 510 are chamfered at 45°. The inflow and outflow conduits20 and 30 are adapted for ease of insertion, subcutaneous tunneling, andremoval, while also providing a resistance to infection and thrombosis.

In one embodiment, a portion of the inflow conduit 20 may be insertedinto the lumen of a blood vessel and advanced to the desired positionusing a percutaneous approach or an open surgical approach. To aid inthe positioning of the inflow and outflow conduits 20 and 30, theconduits may have radiopaque marker bands or other radiopaque materialsembedded within the walls 512 and 514 of the inflow and outflowconduits, respectively, that are visible under fluoroscopy. For example,portions of the inflow and outflow conduits 20 and 30 may be composed ofCarbothane® PC-3575 polyurethane embedded with barium sulfate salts. Inother embodiments the portions of the inflow and outflow conduits 20 and30 that are configured to be inserted into the lumen of the vascularsystem may have self-expanding or radially expansive (such as can beaccomplished by incorporating nitinol) walls so that the diameter of theintravascular portion of the inflow and outflow conduits 20 and 30 willmatch the diameter of the vascular system at that location, such as isseen with the self expanding segment of the GORE® Hybrid Vascular Graft.

In various embodiments, including the embodiment shown in FIG. 37, theinflow and outflow conduits 20 and 30 may be attached to blood vesselsusing a surgical anastomosis, using suture in a running or dividedfashion, henceforth described as an “anastomotic connection.” Ananastomotic connection can also be made with surgical clips and otherstandard ways of making an anastomosis. For example, an anastomoticconnection may be made between the ePTFE distal segment 502 of theoutflow conduit 30 and a blood vessel.

In certain embodiments where an anastomotic connection is made, theoutflow conduit 30 is secured to blood vessels having an initialdiameter between 1 mm and 20 mm, and preferably vessels having aninitial diameter between 1 mm and 6 mm.

Conversely, in other embodiments shown in FIGS. 32A-B and 37-40,portions of the inflow and outflow conduits 20 and 30 are placed withina blood vessel or the right atrium. For example, the distal end 506 ofthe inflow conduit 20 may be positioned within the right atrium or thesuperior vena cava. As shown in FIGS. 32A-32B, the side holes 508 aid inthe aspiration or discharge of blood when the distal end 506 has beenplaced intravascularly.

In various other embodiments, at least one of the inflow and outflowconduits 20 and 30 may be compatible for use with a hemodialysismachine. For example, a patient using the blood pump system 10 may alsoneed to receive a hemodialysis treatment. In this example, blood may bewithdrawn from the blood pump system, passed through a hemodialysismachine, and then discharged back into the blood pump system fordelivery back into the vascular system, thereby eliminating the need tocreate an additional vascular access site in the patient.

As shown in FIG. 35, one embodiment of the control system 14 includes acontrol device 21 having at least one processor 24 and memory 27 fordelivering power to the pump and receiving information from the bloodpump 25, whereby the information is used to set and control pump speedand estimate the flow rate of fluid through the pump system. Theprocessor 24 is configured to read, process, and execute systems,methods, and instructions encoded on a computer-readable medium. Thecontrol system 14 then estimates the wall shear stress in the targetvessel using the measured or estimated vessel diameter and the measuredor estimated average flow rate of the pump system. The control devicealso includes a power source 26, optionally having a battery 28.

In one embodiment, the control system 14 receives sensor feedback fromone or more sensors 122. Any of a variety of suitable sensors may beused to detect any of a variety of changes in a physical quantity of theblood, blood pump 15, the blood pump system 10, and/or the targetvessel. The sensors 122 generate a signal indicative of the change to beanalyzed and/or processed. Essentially, the sensors 122 monitor avariety of properties of the blood pump system 10, the blood flowingthrough the system, and the target blood vessel for changes that can beprocessed and compared to desired reference values or predeterminedstandards. The desired reference values or predetermined standards maybe stored in a database or other suitable medium.

In various embodiments, one or more sensors 122 may be in communicationwith the blood pump 25, the inflow conduit 20, the outflow conduit 30,the donating vessel or location, or the accepting vessel or location. Invarious embodiments, the control system 14 or portions thereof may belocated internally within the housing or casing of the blood pump 25.For example, one or more of the sensors 122 may be located in the inlet110 or outlet 115 of the blood pump 25. In other embodiments, thecontrol system 14 may be external to the pump.

Wall shear stress can be used as a variable to configure the operationof the pump system 10 to result in an increase in the overall diameterand lumen diameter of the target vessel or an increase in the length ofthe target vessel.

Assuming Hagen-Poiseuille blood flow (i.e. laminar flow with a fullydeveloped parabolic velocity profile) in the lumen of a vessel having acircular cross section, then WSS can be determined using the equation:

WSS(Pa)=4Qμ/πR ³   [Eqn. 1]

where:

-   -   Q=flow rate (m³/s)    -   μ=viscosity of blood (Pa/s)    -   R=radius of vessel (m)

Wall Shear Stress Control Method #1: Manual

Mean and/or peak WSS in the target blood vessel can be controlled byadjusting pump speed, which affects the blood flow rate through thepump-conduit system and therefore blood flow through the target vessel.As shown in FIG. 36A, a manual control method 600 may involve the directmeasurement of blood viscosity at block 602 (by sampling the patient'sblood and analyzing it in a viscometer), blood flow rate in the bloodpump system or blood flow rate in the target vessel at block 604 (byplacement of an ultrasonic flow sensor on either the inflow or outflowconduit or by ultrasound or thermal dilution methods, respectively) andvessel radius at block 606 (by various imaging methods includingangiography, ultrasound, computed tomography, or magnetic resonanceimaging). The WSS acting on the vessel wall is determined at block 608,compared to the desired level at blocks 610 or 612, and then the pumpflow rate (Q) is adjusted through changes in the rotational speed of thepump impeller at blocks 614 or 616. Changes in pump speed are effectedby varying the duty-cycle of the pulse width modulation of the motorinput voltage.

Wall Shear Stress Control Method #2: Automatic With Indirect BloodViscosity, Direct Blood Flow, and Target Blood Vessel DiameterMeasurements

An automatic WSS control system may involve direct measurement of bloodflow rate in the pump system or the target vessel, and directmeasurement of the diameter of the target vessel blood vessel. As shownin FIG. 36B, this automatic WSS control method 620 may involve indirectmeasurements of blood viscosity at block 622 (estimated based on itsknown relationship with measured hematocrit and approximate mean WSS).Periodic calibration of the viscosity estimator at block 624 may beperformed using direct measurements of viscosity as previouslydescribed. In clinical practice, the blood viscosity usually variesslowly.

Wall Shear Stress Control Method #3: Automatic With Indirect BloodViscosity, Blood Flow, Target Blood Vessel Diameter Measurements, andDirect Vein Pressure Measurements

As shown in FIG. 36C, an automatic WSS control method 630 may involveindirect measurements of blood viscosity (estimated based on its knownrelationship with measured hematocrit and approximate mean WSS) at block622, blood flow rate through the blood pump system (estimated based onits relationship to motor state variables) at block 632, measurements ofthe target blood vessel pressure at block 634, and measurements of thevessel radius (estimated based on vascular resistance) at block 638.Vascular resistance is calculated at block 636 based on the estimatedpump flow rate and the measured blood pressure in the vessel. Periodiccalibration of the blood viscosity, pump flow, and target vessel radiusestimators respectively, may be performed using direct measurements atblocks 624, 640, and 642, respectively, as previously described.

Wall Shear Stress Control Method #4: Automatic With Indirect BloodViscosity, Blood Flow, Pump Pressure Head, and Target Blood VesselDiameter Measurements

As shown in FIG. 36D, an automatic WSS control method 650 may involveindirect measurements of blood viscosity (estimated based on its knownrelationship with measured hematocrit and approximate mean WSS) at block622, blood flow rate through the blood pump system (estimated based onits relationship to motor state variables) at block 632, and vesselradius (estimated based on vascular resistance) at block 638. Vascularresistance is calculated at block 636 based on the pump flow rateestimated at block 632 and pump pressure head, where pump pressure headis also estimated at block 652 based on its relationship to motor statevariables. Periodic calibration of the blood viscosity, pump flow, andtarget vessel radius estimators may be performed using directmeasurements at blocks 624, 640, and 642, respectively, as previouslydescribed. Periodic calibration of the pump pressure head estimator maybe performed by measuring pump inlet and pump outlet pressures withseparate pressure transducers and calculating their difference at block654, or by directly measuring pressure head across the pump with adifferential pressure sensor.

Sensorless Determination of Blood Pump System Flow Rate and PressureHead:

Referring to FIG. 35, the processor 24 is adapted to detect and monitorelectric current appearing in one or more of the electric coils of thecoil assembly 170 of the pump via the power cable 120 which, inconjunction with monitoring the voltage provided to the coil assemblypermits the processor 24 to derive the input power (P_(in)) consumed bythe blood pump 25 and an actual rotational speed of the impeller 140(ω). The processor 24 can estimate pump flow rate (Q) or changes in flowrate (ΔQ) as a function of P_(in) and ω. For example, Q=f[P_(in), ω].More specifically, the following equation is used:

Q=a+b·ln(P _(in))+c·ω ^(0.5)   [Eqn. 2]

where:

-   -   Q=flow rate (L/min)    -   P_(in)=Motor input power (W)    -   ω=Pump speed (rpm)        Motor input power is derived from the measured motor current and        voltage. The values for a, b, and c are derived from curve        fitting the plot of pump flow rate as a function of motor speed        and input power.

The processor 24 can also estimate pump pressure head (H_(p)) or changesin pump pressure head (ΔH_(p)) as a function of P_(in) and ω. Forexample, H_(p)=f[P_(in), ω]. More specifically, the following equationis used:

H _(p) =d+e·ln(P _(in))+f·ω ^(2.5)   [Eqn. 3]

The values for d, e, and f are derived from curve fitting the plot ofpump pressure head as a function of pump speed and motor input power,where H_(p) is measured across the inflow conduit 20, pump 25, andoutflow conduit 30.

Determination of Vascular Resistance and Estimation of Vessel Radius:

Vascular resistance (Rv) is the resistance to flow that must be overcometo push blood through the circulatory system. Resistance is equal todriving pressure (H_(v)) divided by the flow rate. When the blood pumpsystem is connected to a target vessel that is a vein, the vascularresistance is calculated using the following equation:

R _(v)=(P _(v) −CVP)/Q   [Eqn. 4]

where:

-   -   H_(v)=pressure head lost across the peripheral vessel on the        return path of the blood to the heart (mmHg)    -   P_(v)=vein pressure at anastomosis (mmHg)    -   CVP=central venous pressure (mmHg)    -   R_(v)=vascular resistance ((mmHg·min)/L)        Normally, CVP ranges between 2-8 mmHg and can be neglected in        the above equation because the operating ranges of P_(v) and Q        are proportionally much greater. As illustrated in FIG. 36E,        vascular resistance can be represented graphically as the slope        of various P_(v) vs. Q curves 660. Since the curves 660 are        nonlinear, the slope is a function of Q. As illustrated by the        following equation, the vascular resistance may be derived by        temporarily increasing speed by several hundred rpm (Δω),        measuring the resulting change in vein pressure (ΔP_(v)), and        estimating the resulting change in pump flow (ΔQ):

R _(v)(Q)=ΔP _(v) /ΔQ   [Eqn. 5]

It is noted that the vascular resistance is a strong function of vesseldiameter or radius, with smaller veins having high vascular resistance.Vascular resistance can be quantified in various units, for example,Wood units ((mmHg·min)/L) can be multiplied by eight to convert to SIunits ((Pa·s)/m³).

Alternatively, pump pressure head (H_(p)) may be used as a basis forcalculating vascular resistance. When the pump-conduit system isconfigured to withdraw blood from one location in the vascular system todischarge it into a peripheral artery or vein it is a reasonableassumption that the pressure head gained across the system (Hp) isexactly equal to the pressure head lost across the peripheral vessel onthe return path of the blood to the heart (H_(v)):

H_(v)=H_(p)   [Eqn. 6]

The radius of the peripheral vessel is inversely proportional to itsvascular resistance (R_(v)), the ratio of H_(v) to Q. AssumingHagen-Poiseuille blood flow in the vessel of circular cross section, thevascular resistance can be represented using the equation:

R _(v)(Pa·s/m³)=P _(v) /Q=8·μ·L/π·R ⁴   [Eqn. 7]

where:

-   -   P_(v) is expressed in units of Pa    -   Q is expressed in units of (m³/s)    -   μ=viscosity of blood (Pa/s)    -   R=radius of vessel (m)    -   L=length of vessel (m)        In practice, Eqn. 7 would be refined based upon in vivo        measurements of pressure drop across specific veins of known        diameter. This provides an empirical form of the equation:

R _(v)(Pa·s/m³)=K·μ/R ⁴   [Eqn. 8]

where:

-   -   K is an empirical constant for the target vein (m)

Determination of Wall Shear Stress:

The wall shear stress in the target vessel can be determined based onthe above equations. Using Eqn. 4, the pump flow rate can be expressedaccording to the following equation:

Q=P _(v) /R _(v)   [Eqn. 9]

Using Eqn. 8, vessel radius can be expressed according to the followingequation:

R=(K·μ/R _(v))^(0.25)   [Eqn. 10]

Using Eqns. 1, 9, and 10, the wall shear stress can be expressedaccording to the following equation:

WSS(Pa)=((4·P _(v))/(π·K ^(0.75)))·(μ/R _(v))^(0.25)   [Eqn. 11]

In various embodiments, the estimated variables used by the controlsystem are periodically calibrated. For example, the estimates of flowrate and pressure head are periodically calibrated using actual measuredvalues at an interval ranging from 1 minute and up to 30 days.Similarly, the estimate of artery or vein radius is periodicallycalibrated using actual measured values at an interval ranging from 1minute and up to 30 days.

Safety Features and Alarms:

The automatic control system may also include safety features to avoidhazards associated with changes in the patient's cardiovascular systemor malfunctions of the pump system or pump control system. As shown inFIG. 36F, a speed control method 670 can detect characteristic changesin the motor current waveform associated with decreased preload orincrease in afterload (e.g. due to thrombosis), suction, flowlimitation, and imminent collapse of the vessel around the inflowconduit tip at block 672. Spectral analysis of the motor currentwaveform is performed using a Fourier transform at block 674. When theamplitude of the second harmonic term of the Fourier series exceeds apredetermined value at block 676 , suction has occurred and collapse isdeemed imminent. Pump speed is immediately decreased at block 616 and analarm is triggered at block 678A within the control device 21. Whennormal operation is restored, the alarm is canceled at block 678B.

As shown in FIG. 36G, a speed control method 680 can detect low flowconditions. When the pump flow rate drops below the safe threshold levelto avoid thrombosis of the pump-conduit system 10 at block 682, the pumpspeed is immediately increased at block 614 and an alarm is triggered atblock 678A within the control device 21. When normal operation isrestored, the alarm is canceled at block 678B.

As shown in FIG. 36H, a speed control method 690 can detect high wallshear stress conditions. When the WSS rises above the safe thresholdlevel to avoid damage to the vessel endothelium at block 692, the pumpspeed is immediately decreased at block 616 and an alarm is triggered atblock 678A within the control device 21. When normal operation isrestored, the alarm is canceled at block 678B.

In yet another embodiment in which the inflow conduit 20 is connected toan artery and the outflow conduit 30 is connected to a vein, the controlsystem 14 monitors and modifies the pulsatility of blood flow that isdischarged into the accepting vein. For example, the control system 14can monitor the electrocardiogram or monitor the cyclic changes in thepulse wave of blood coming into the blood pump system. Duringventricular contraction and pulse wave propagation, the control systemcan decrease the rotational speed of the pump. During systole and afterthe pulse wave has passed, the control system can increase therotational speed of the pump. In this manner, pulsatility in the bloodentering the accepting vein can be reduced. Alternatively, thepulsatility of the blood in the accepting vein may be periodicallychecked manually, as may be accomplished with ultrasound, and the pumpmay be manually adjusted, for example, by tuning the head-flowcharacteristics of the pump, adding a compliance reservoir or elasticreservoir (a segmental or a diffuse change) to the pump inflow oroutflow, or modulating the pump speed. Other adjustments may also bemade. Alternatively, a compliance reservoir or elastic reservoir can beadded to the inflow or outflow conduits at the time of implantation ofthe blood pump system.

In various other embodiments, the control system 14 is monitored andadjusted manually or with a software program or application encoded on acomputer-readable medium and executable by the processor 24, or otherautomated systems. The computer-readable medium may include volatilemedia, nonvolatile media, removable media, non-removable media, and/oranother available medium that can be accessed by control system 14. Byway of example and not limitation, the computer-readable medium mayinclude computer storage media and communication media. Computer storagemedia includes memory, volatile media, nonvolatile media, removablemedia, and/or non-removable media implemented in a method or technologyfor storage of information, such as computer readable instructions, datastructures, program modules, or other data.

The software program may include executable instructions toautomatically adjust the pump speed to maintain the desired amount ofblood flow, mean blood speed or velocity, and mean WSS in the vesselsegment to be treated (the “target vessel” or the “target blood vessel”)in which a persistent increase in overall diameter and lumen diameter,or length, is desired, whether it is a donating artery, a donating vein,an accepting artery, or an accepting vein. Alternatively, the overalldiameter, lumen diameter, length, and blood flow in the target vesselmay be periodically checked manually, as may be accomplished withultrasound, and the pump may be manually adjusted, for example, bytuning the head-flow characteristics of the pump or modulating the pumpspeed. Other adjustments may also be made.

In one embodiment, the mean blood speed is determined by calculating anaverage of multiple discrete measurements of blood speed by summing thediscrete measurements and dividing the total by the number ofmeasurements. Mean blood speed can be calculated by taking measurementsover a period of milliseconds, seconds, 1 minute, 5 minutes, 15 minutes,30 minutes, 1 hour, or multiple hours.

In another embodiment, the mean WSS is determined by making a series ofdiscrete measurements, making multiple discrete determinations of WSS(using those measurements), summing the discrete WSS determinations, anddividing the total by the number of determinations. Mean WSS can becalculated by taking measurements and making discrete WSS determinationsover a period of seconds, 1 minute, 5 minutes, 15 minutes, 30 minutes, 1hour, or multiple hours.

In one embodiment, the control system 14 receives information fromsensor 22 in communication with the blood pump 25. In other embodiments,the control system 14 receives information from a sensor 22 incommunication with an inflow conduit 20 or an outflow conduit 30 or in avessel in fluid communication the inflow or outflow conduit. In variousembodiments, all or portions of the control system 14 may be locatedwithin the pump body 25, while in other embodiments all or a portion ofthe control system may be located within the conduits, or within thecontrol device 21.

The systems and methods described herein increase the mean WSS level inperipheral veins and arteries. Normal mean WSS for veins ranges between0.076 Pa and 0.76 Pa. The systems described herein are configured toincrease the mean WSS level in the accepting peripheral vein to a rangebetween 0.76 Pa and 23 Pa, preferably to a range between 2.5 Pa and 10Pa. Normal mean WSS for arteries ranges between 0.3 Pa and 1.5 Pa. Forartery dilation, the systems and methods described herein increase themean WSS level to a range between 1.5 Pa and 23 Pa, preferably to arange between 2.5 Pa and 10 Pa. In certain instances, sustained mean WSSless than 0.76 Pa in veins or less than 1.5 Pa in arteries may increasethe overall diameter and lumen diameter of these vessels but the extentand rate of this increase is not likely to be clinically meaningful orcompatible with routine clinical practice. Sustained mean WSS greaterthan 23 Pa in arteries or veins is likely to cause denudation (loss) ofthe endothelium of the blood vessels, or damage to the endothelium,which is known to retard dilation of blood vessels in response toincreases in mean blood speed and mean WSS. Pumping blood in a mannerthat increases mean WSS to the desired range for preferably 1 day to 84days, and more preferably between about 7 and 42 days, for example,produces a persistent increase in the overall diameter and lumendiameter in an accepting vein, a donating vein, or a donating arterysuch that veins and arteries that were initially ineligible orsuboptimal for use as a hemodialysis access sites or bypass grafts dueto small vein or artery diameter become usable or more optimal. Theblood pumping process may be monitored and adjusted periodically. Forexample, the pump may be adjusted over a period of minutes, hours, 1day, 3 days, 1 week, or multiple weeks to account for changes in theperipheral vein or artery (such as a persistent increase in the overalldiameter and lumen diameter) prior to achieving the desired persistentdilation.

Referring to FIGS. 37-40, a system 10 to increase the overall diameterand lumen diameter of veins and arteries is illustrated as used for apatient 1. In FIG. 37, the system 10 draws deoxygenated venous bloodfrom the patient's venous system and discharges that blood into theaccepting peripheral vessel 700. The system 10 also increases the meanspeed of blood in the accepting peripheral vessel 700 and increases themean WSS exerted on the endothelium of the accepting peripheral vessel700, to increase the overall diameter and lumen diameter of theaccepting peripheral vessel 700 located, for example, in an arm or leg.The diameter of blood vessels such as peripheral veins can be determinedby measuring the diameter of the lumen, which is the open space at thecenter of blood vessel where blood is flowing or by measuring thediameter of the overall vessel, which includes the open space and thewalls of the blood vessel.

The invention also relates to simultaneously and persistently increasingthe overall diameter and lumen diameter of a peripheral vein or arteryby directing blood into or out of the peripheral vein or artery, therebyincreasing the mean speed of the blood in the peripheral vein or arteryand increasing the mean WSS on the endothelium of the peripheral vein orartery. Systems are described wherein the mean speed of the blood in aperipheral vein or artery and the mean WSS on the endothelium of theperipheral vein or artery is increased by using a blood pump system.Preferably, the pump directs blood into the peripheral vein, wherein thepumped blood has reduced pulsatility, such as when the pulse pressure islower than blood in a peripheral artery.

The system 10 is suitable to maintain a flow rate preferably between 50mL/min and 2500 mL/min and optionally between 50 mL/min and 1000 mL/minwhile also maintaining a pressure range between 25 mmHg and 350 mmHg. Aspreviously described, the control system 14 may be optimized to maintaina steady mean wall shear stress of between 0.76 Pa and 23 Pa inperipheral veins such that the overall diameter and lumen diameter ofthe peripheral veins are persistently increased by as much as 5% to morethan 200%.

The systems described herein also increase the mean speed of blood inperipheral veins. At rest, the mean speed of blood in the cephalic veinin humans is generally between 5 to 9 cm/s (0.05 to 0.09 m/s). For thesystems described herein, the mean speed of blood in the peripheral veinis increased to a range between 10 cm/s and 120 cm/s (0.1 and 1.2 m/s),preferably to a range between 25 cm/s and 100 cm/s (0.25 m/s and 1.0m/s), depending on the initial overall diameter or lumen diameter ofperipheral accepting vein and the final overall or lumen diameter thatis desired. The systems described herein also increase the mean speed ofblood in peripheral arteries. At rest, the mean speed of blood in thebrachial artery is generally between 10 and 15 cm/s (0.1 and 0.15 m/s).For the systems and methods described herein, the mean speed of blood inthe peripheral artery is increased to a range between 10 cm/s and 120cm/s (0.1 and 1.2 m/s), preferably to a range between 25 cm/s and 100cm/s (0.25 and 1.0 m/s), depending on the initial overall diameter orlumen diameter of artery the final overall or lumen diameter that isdesired.

Preferably, the mean blood velocity is increased for between 1 day and84 days, or preferably, between 7 and 42 days, to induce a persistentincrease in the overall diameter and lumen diameter in the peripheralaccepting vein, peripheral accepting artery, peripheral donating vein,or peripheral donating artery such that veins and arteries that wereinitially ineligible or suboptimal for use as a hemodialysis access siteor bypass graft due to a small vein or artery diameter become usable.This can also be accomplished by intermittently increasing mean bloodvelocity during the treatment period, with intervening periods of normalmean blood velocity.

Studies have shown that baseline hemodynamic forces and changes inhemodynamic forces within veins and arteries play a vital role indetermining the overall diameter and lumen diameter, and the length ofthose veins and arteries. For example, persistent increases in meanblood velocity and mean WSS can lead to a persistent increase in thelumen diameter and overall diameter, and length, of veins and arteries.The elevated mean blood velocity and mean WSS are sensed by endothelialcells, which trigger signaling mechanisms that result in stimulation ofvascular smooth muscle cells, attraction of monocytes and macrophages,and synthesis and release of proteases capable of degrading componentsof the extracellular matrix such as collagen and elastin. As such, thepresent invention relates to increasing mean blood velocity and mean WSSfor a period of time sufficient to result in vein and artery remodelingand an increase in the overall diameter and the lumen diameter, andlength, of the veins and arteries.

The systems described herein increase the mean WSS level in a peripheralvein or artery. Normal mean WSS for veins ranges between 0.076 Pa and0.76 Pa. The systems described herein increase the mean WSS level inveins to a range between 0.76 Pa and 23 Pa, preferably to a rangebetween 2.5 Pa and 10 Pa. Normal mean WSS for arteries ranges between0.3 Pa and 1.5 Pa. To persistently increase the overall diameter andlumen diameter of arteries, the systems and methods described hereinincrease the mean WSS level to a range between 1.5 Pa and 23 Pa,preferably to a range between 2.5 Pa and 10 Pa. Preferably, the mean WSSis increased for between 1 days and 84 days, or preferably, between 7and 42 days, to induce a persistent increase in the overall diameter andlumen diameter in the peripheral accepting vein, peripheral acceptingartery, peripheral donating vein, or peripheral donating artery suchthat veins and arteries that were initially ineligible or suboptimal foruse as a hemodialysis access site or bypass graft due to a small veinand artery diameter become usable. This can also be accomplished byintermittently increasing mean WSS during the treatment period, withintervening periods of normal mean WSS.

In some circumstances, sustained periods of mean WSS levels in theperipheral veins lower than 0.076 Pa or in peripheral arteries lowerthan 1.5 Pa may result in increased overall diameter and lumen diameterof these veins and arteries, but the extent and rate of this increase isnot likely to be clinically meaningful or compatible with routineclinical practice. Sustained mean WSS levels in peripheral veins andarteries higher than about 23 Pa are likely to cause denudation (loss)of the endothelium of the veins or damage to the endothelium of theveins. Denudation of the endothelium or damage to the endothelium ofblood vessels is known to reduce the increase in overall diameter andlumen diameter of blood vessels in the setting of increased in meanblood velocity and mean WSS. The increased mean WSS induces sufficientpersistent increase in the overall diameter and lumen diameter, orlength, in the veins and arteries, such that those that were initiallyineligible or suboptimal for use as a hemodialysis access site or bypassgraft due to a small vein or artery diameter become usable or moreoptimal. The diameter of the peripheral accepting vein, peripheralaccepting artery, peripheral donating vein, or peripheral donatingartery can be determined intermittently, such as every 1 day, 3 days, 1week, or multiple weeks for example, to allow for pump speed adjustmentin order to optimize the rate and extent of the persistent increase inthe overall diameter and lumen diameter of the vein and artery duringthe treatment period.

The systems described herein also increase the mean speed of blood inperipheral veins. At rest, the mean speed of blood in the cephalic veinin humans is generally between 5 and 9 cm/s (0.05 and 0.09 m/s). For thesystems described herein, the mean speed of blood in the peripheral veinis increased to a range between 10 cm/s and 120 cm/s (0.1 and 1.2 m/s),preferably to a range between 25 cm/s and 100 cm/s (0.25 m/s and 1.0m/s), depending on the initial overall diameter or lumen diameter of theperipheral accepting vein and the desired final overall diameter andlumen diameter of the peripheral accepting vein. The systems describedherein also increase the mean speed of blood in peripheral arteries. Atrest, the mean speed of blood in the brachial artery is generallybetween 10-15 cm/s (0.1 and 0.15 m/s). For the systems and methodsdescribed herein, the mean speed of blood in the peripheral artery isincreased to a range between 10 cm/s and 120 cm/s (0.1 and 1.2 m/s),preferably to a range between 25 cm/s and 100 cm/s (0.25 and 1.0 m/s),depending on the initial overall diameter or lumen diameter of theperipheral artery and the desired final overall diameter or lumendiameter of the peripheral artery. Preferably, the mean blood velocityis increased for between 1 day and 84 days, or preferably, between 7 and42 days, to induce a persistent increase in the overall diameter and thelumen diameter, or length, of the peripheral accepting vein, peripheralaccepting artery, peripheral donating vein, or peripheral donatingartery such that veins and arteries that were initially ineligible orsuboptimal for use as a hemodialysis access site or bypass graft due toa small vein or artery diameter or inadequate length become usable. Meanblood velocity levels in the accepting peripheral vein, peripheralaccepting artery, peripheral donating vein, or peripheral donatingartery lower than 10 cm/s (0.1 m/s) may result in increased overalldiameter and lumen diameter of these veins and arteries, but the extentand rate of this increase is not likely to be clinically meaningful orcompatible with routine clinical practice. Mean blood velocity levels inperipheral accepting veins, peripheral accepting arteries, peripheraldonating veins, or peripheral donating arteries higher than about 120cm/s (1.2 m/s) are likely to cause denudation (loss) of the endotheliumof the veins or damage to the endothelium of veins. Denudation or damageof the endothelium of blood vessels is known to reduce the increase inthe overall diameter and lumen diameter of blood vessels observed in thesetting of increased mean blood velocity. The increased mean bloodvelocity in the desired range and for a sufficient period of timeinduces sufficient persistent increase in the overall diameter and lumendiameter, or length, in the veins and arteries, such that those thatwere initially ineligible or suboptimal for use as a hemodialysis accesssite or bypass graft due to a small vein or artery diameter orinadequate length become usable. The overall diameter or lumen diameterof the peripheral accepting vein, peripheral accepting artery,peripheral donating vein, and peripheral donating artery can bedetermined intermittently, such as every minute(s), hour(s), 1 day, 3days, 1 week, or multiple weeks for example, to allow for pump speedadjustment in order to optimize the rate and extent of the persistentincrease in the overall diameter and lumen diameter of the vein andartery during the treatment period.

In one embodiment shown in FIG. 34, the system 10 includes the bloodpump 25, the pair of conduits 12, and the control device 21 for movingdeoxygenated venous blood from a donating vein or location in the venoussystem of a patient to a peripheral accepting vein. In variousembodiments, the peripheral accepting vein may be a cephalic vein,radial vein, median vein, ulnar vein, antecubital vein, median cephalicvein, median basilic vein, basilic vein, brachial vein, lesser saphenousvein, greater saphenous vein, femoral vein, or other veins. Other veinsthat might be useful in the creation of a hemodialysis access site orbypass graft or other veins useful for other vascular surgery proceduresrequiring the use of veins may be used. The conduits 12 move thedeoxygenated blood to the peripheral accepting vein. The persistentlyelevated mean speed of the blood and the elevated mean WSS in theperipheral vessel causes a persistent and progressive increase in theoverall diameter and lumen diameter of the peripheral accepting vein.Thus, the system 10 of the present invention advantageously increasesthe diameter or length of the peripheral vein 4 so that it can be used,for example, to construct an hemodialysis access site (such as an AVF orAVG), a bypass graft, or used in another clinical setting where a veinof a certain diameter or length is needed, as determined by one skilledin the art.

As used herein, deoxygenated blood is blood that has passed through thecapillary system and had oxygen removed by the surrounding tissues andthen passed into the venous system. A peripheral vein, as used herein,means any vein with a portion residing outside of the chest, abdomen, orpelvis. In the embodiment shown in FIG. 37, the peripheral acceptingvein 712 is the cephalic vein. However, in other embodiments, theperipheral accepting vein may be a radial vein, median vein, ulnar vein,antecubital vein, median cephalic vein, median basilic vein, basilicvein, brachial vein, lesser saphenous vein, greater saphenous vein,femoral vein, or other veins. In addition to a peripheral vein, otherveins that might be useful in the creation of a hemodialysis access siteor bypass graft or other veins useful for other vascular surgeryprocedures requiring the use of veins may also be used as acceptingveins, such as those residing in the chest, abdomen, and pelvis.

FIG. 37 illustrates another embodiment for using the system 10 toincrease the overall diameter and lumen diameter of a blood vessel. Inthis embodiment, the system 10 is configured to remove deoxygenatedblood from a donating vein 700 and move the blood to the superior venacava or right atrium 702 of the heart 704. As shown, an inflow conduit706 is connected in fluid communication with the donating vein 700, inthis case the cephalic vein. In one embodiment, the connection may bemade using a short ePTFE segment of the inflow conduit 706 that is usedto secure the inflow conduit 706 to the donating vein 700 while theremaining segment of the inflow conduit is made using polyurethane. Inother embodiments, at least a portion of the inflow conduit or theoutflow conduit further comprises nitinol, for kink and compressionresistance. As shown, one end of the outflow conduit 710 is connected tothe blood pump 25 while the other end of the outflow conduit is fluidlyconnected to the superior vena cava and the right atrium 702 by anintravascular portion. For the embodiment of FIG. 37, a blood pump isused increase the rate at which blood moves from the donating vein 700to the superior vena cava and right atrium 702 of the heart 704 in orderto achieve a desired elevated level of mean blood velocity and elevatedlevel of mean WSS in the donating vein 700. The pump is operated at arate and for a time sufficient to result in a desired persistentincrease in the overall diameter and lumen diameter of the donatingvein, such as a 10% increase, a 25% increase, a 50% increase, or anincrease of 100% or more from the starting diameter. In a furtherembodiment, one or more venous valves between the junction of the inflowconduit 706 and the donating vein 700, and the right atrium 702 may berendered incompetent or less competent (using any of the methodsavailable to one skilled in the art) to allow blood to flow in aretrograde fashion in the donating vein 700 and then into the inflowconduit 706.

FIG. 38 illustrates another embodiment for using the system 10 toincrease the overall diameter and lumen diameter of a blood vessel. Inthis embodiment, the system 10 is configured to remove oxygenated bloodfrom a donating artery 712 (in this case the brachial artery) and movethe blood to the superior vena cava and right atrium 702 of the heart704. As shown, an inflow conduit 706 is connected in fluid communicationwith the donating artery 712. In one embodiment, the connection may bemade using a short ePTFE segment of the inflow conduit 706 that is usedto secure the inflow conduit to the donating artery 712 while theremaining segment of the inflow conduit is made using polyurethane. Inother embodiments, one or both segments of the inflow conduit 706further comprise nitinol, such as for kink and compression resistance.As shown, one end of the outflow conduit 710 is connected to the bloodpump 25 while the other end of the outflow conduit is fluidly connectedto the superior vena cava and the right atrium 702 by an intravascularportion. For the embodiment of FIG. 38, a blood pump is used increasethe rate at which blood moves from the donating artery 712 to the rightatrium 702 of the heart 704 in order to achieve a desired elevated levelof mean blood velocity and elevated mean level of WSS in the donatingartery 712. The pump is operated at a rate and for a time sufficient toresult in a desired persistent increase in the overall diameter andlumen diameter of the donating artery, such as a 10% increase, a 25%increase, a 50% increase, or an increase of 100% or more from thestarting diameter.

In other embodiments, oxygenated arterial blood may be moved from adonating artery to an accepting location. Donating arteries may include,but are not limited to, a radial artery, ulnar artery, interosseousartery, brachial artery, anterior tibial artery, posterior tibialartery, peroneal artery, popliteal artery, profunda artery, superficialfemoral artery, or femoral artery.

FIG. 39 illustrates another embodiment for using the system 10 toincrease the overall diameter and lumen diameter of a blood vessel. Inthis embodiment, the system 10 is configured to remove oxygenated bloodfrom a donating artery 712 (in this case the brachial artery) and movethe blood to the superior vena cava and right atrium 702 of the heart704. As shown, a conduit 716 is connected in fluid communication withthe donating artery 712. In one embodiment, the connection may be madeusing a short ePTFE segment of the conduit 716 that is used to securethe inflow conduit to the donating artery 712 while the remainingsegment of the inflow conduit is made using polyurethane. In otherembodiments, one or both segments of the conduit 716 further comprisenitinol, such as for kink and compression resistance. For the embodimentof FIG. 39, there is no pump and blood moves passively from the higherpressure donating artery 712 to the lower pressure superior vena cavaand right atrium 702, and the conduit 716 is configured in length andlumen diameter to achieve a desired elevated level of mean bloodvelocity and mean WSS in the donating artery 712. The conduit 716remains in place for a time sufficient to result in a desired persistentincrease in the overall diameter and lumen diameter of the donatingartery 712, such as a 10% increase, a 25% increase, a 50% increase, oran increase of 100% or more from the starting diameter.

FIG. 40 illustrates another embodiment for using the system 10 toincrease the overall diameter and lumen diameter of a peripheral artery.In this embodiment, the system 10 is configured to remove oxygenatedblood from a target artery 718, such as the radial artery, and move theblood to an accepting artery 720, such as the brachial artery. As shown,an inflow conduit 706 is connected in fluid communication with thetarget artery 718. In one embodiment, the connection between the inflowconduit 706 and an artery or the outflow conduit 710 and an artery maybe made using a short ePTFE segment of the respective conduit that isused to fluidly connect the inflow conduit to the target artery 718 orthe outflow conduit 710 that is fluidly connected to the acceptingartery 720, while the remaining segments of the inflow and outflowconduits can be made using polyurethane. In other embodiments, one orboth segments of the inflow conduit 706 or the outflow conduit 710further comprise nitinol, such as for kink and compression resistance.

As shown, one end of the outflow conduit 710 is connected to the bloodpump 25 while the other end of the outflow conduit is fluidly connectedto the accepting artery 720. For the embodiment of FIG. 40, the bloodpump 25 is used increase the rate at which blood is withdrawn from thetarget artery 718 in order to achieve a desired elevated level of meanblood velocity and elevated mean level of WSS in the target artery. Thepump is operated at a rate and for a time sufficient to result in adesired persistent increase in the overall diameter and lumen diameterof the target artery 718, such as a 10% increase, a 25% increase, a 50%increase, or an increase of 100% or more from the starting diameter.

While the invention has been explained in relation to exemplary aspectsand embodiments, it is to be understood that various modificationsthereof will become apparent to those skilled in the art upon readingthe description. Therefore, it is to be understood that the inventiondisclosed herein is intended to cover such modifications as fall withinthe scope of the appended claims.

1-171. (canceled)
 172. A blood pump system comprising: a blood pump; oneor more conduits, and, a control system to monitor the blood pump systemand modify the operation of the blood pump to maintain an increased meanwall shear stress within an artery or vein fluidly connected to theblood pump.
 173. The blood pump system of claim 1, wherein the bloodpump is a rotary blood pump.
 174. The blood pump system of claim 1,wherein the blood pump is a centrifugal blood pump.
 175. The blood pumpsystem of claim 1, wherein the system is configured to maintain meanwall shear stress within a vein in the range of 0.76 to 23 Pa, or 2.5 to10 Pa.
 176. The blood pump system of claim 1, wherein the system isconfigured to maintain mean blood speed within an artery or vein in therange of 10 cm/s and 120 cm/s, or 25 cm/s and 100 cm/s.
 177. The bloodpump system of claim 1, wherein the one or more conduits have a combinedlength between 4 cm and 220 cm.
 178. The blood pump system of claim 1,wherein the one or more conduits have an inner diameter between 2 mm and10 mm.
 179. The blood pump system of claim 1, wherein at least a segmentof the one or more conduits comprises an elastic reservoir.
 180. Theblood pump system of claim 1, wherein at least a portion of the one ormore conduits comprises a shape-memory alloy, a self-expanding material,or a radially expansive material.
 181. The blood pump system of claim 9,wherein the shape-memory alloy is nitinol.
 182. The blood pump system ofclaim 10, wherein the one or more conduits comprises braided nitinol.183. The blood pump system of claim 11, wherein the one or more conduitscomprises coiled nitinol.
 184. The blood pump system of claim 1, whereinan end of the one or more conduits that are configured to fluidlyconnect to the vascular system in a patient are chamfered at an anglebetween 10 degrees and 80 degrees.
 185. The blood pump system of claim1, wherein an end of the one or more conduits that are configured tofluidly connect to the vascular system in a patient comprises aplurality of holes.
 186. The blood pump system of claim 1, wherein theone or more conduits are connected to the blood pump using aradially-compressive connector.
 187. The blood pump system of claim 1,wherein the one or more conduits include a side port to access a fluidpath within the one or more conduits.
 188. A blood pump systemcomprising: a centrifugal blood pump comprising: a pump housing defininga pump inlet comprising an inflow diffuser to receive blood and directblood onto an impeller, the pump housing having a top bezel and toppivot bearing extending from a top of the housing into the inlet, and abottom bezel and bottom pivot bearing extending from a bottom of thehousing into the interior space of the housing; the impeller suspendedwithin the housing, the impeller having: an impeller pivot having afirst end to engage the top pivot and a second end to engage the bottompivot; a plurality of blades on the top surface of the impeller andextending radially away from a center of the impeller, the blades toforce blood received at the inlet through the pump housing and to theoutlet; and, at least one lumen extending parallel to a central axis ofthe impeller from the bottom surface through the impeller to a topsurface; at least one magnet mechanically engaged to the impeller; atleast one inflow conduit and at least one outflow conduit; and, anelectric motor to magnetically engage the at least one magnet, whereinthe electric motor rotates the at least one magnet and the impeller.189. The blood pump of claim 17, wherein at least one of the pluralityof blades is arcuate.
 190. The blood pump system of claim 17 furthercomprising: a control device for controlling pump speed, the controldevice including: a power source; memory; a processor for controllingthe pump speed and analyzing the feedback; and, a cable for electricallyconnecting the control device to the pump.
 191. The blood pump system ofclaim 17, wherein at least one of the inflow conduit and the outflowconduit is connected to the blood pump using a radially-compressiveconnector.
 192. The blood pump system of claim 17, wherein at least oneof the inflow conduit and the outflow conduit includes a side port toaccess a fluid path within the inflow conduit or the outflow conduit.193. A method of determining and controlling a wall shear stress in ablood vessel comprising: a) measuring a blood viscosity; b) measuring ablood flow rate in a blood pump system or the blood vessel; c) measuringa radius of the blood vessel; d) determining the wall shear stress inthe blood vessel from the measured blood viscosity, the measured flowrate, and the radius of the blood vessel; e) comparing the determinedwall shear stress to a predetermined reference value; f) adjusting ablood pump speed when the determined wall shear stress does notapproximate the predetermined reference value; and g) repeating stepsa-f until the determined wall shear stress approximates thepredetermined reference value.
 194. The method of claim 22 where thewall shear stress is periodically determined and adjusted as necessary,at an interval of 1 ms, 10 ms, 100 ms, 1 s, 1 min, 1 hour, 1 day, 7days, 14 days, or 28 days.
 195. The method of claim 22 where the bloodpump speed is increased and an alarm indication is triggered when theblood flow rate drops below a predetermined safety reference value. 196.The method of claim 22 where the blood pump speed is decreased and analarm indication is triggered when the wall shear stress exceeds thepredetermined safety reference value.
 197. A method of estimating andcontrolling a wall shear stress in a blood vessel using a blood pumpsystem comprising: a) estimating a blood viscosity; b) measuring atleast one motor state variable of the blood pump system selected from avoltage, a current, or a pump speed; c) estimating a blood flow rate anda pressure head in the blood pump system; d) calculating a vascularresistance of the blood vessel from the estimated blood flow rate andthe estimated pressure head; e) estimating a radius of the blood vessel;f) determining the wall shear stress from the estimated blood viscosity,the estimated blood flow rate, and the radius of the blood vessel; g)comparing the determined wall shear stress with a predeterminedreference value; h) adjusting the pump speed when the determined wallshear stress does not approximate the predetermined reference value; andi) repeating steps a-h until the determined wall shear stressapproximates the predetermined reference value.
 198. The method of claim26 where the wall shear stress is periodically determined and adjustedas necessary, at an interval of 1 ms, 10 ms, 100 ms, 1 s, 1 min, 1 hour,1 day, 7 days, 14 days, or 28 days.
 199. The method of claim 30 wherethe blood viscosity is periodically estimated based on its knownrelationship with a measured hematocrit and an approximate wall shearstress at an interval of about 1 to 14 days.
 200. A sensorless method toavoid a collapse of a blood vessel fluidly connected to a blood pumpsystem upon detecting an imminence of the collapse at an inlet of theblood pump system comprising: a) measuring a blood pump motor current;b) continually determining a spectral analysis representation of theblood pump motor current in a form of a Fourier series; c) providing adetection indication when an amplitude of the second harmonic term ofthe Fourier series exceeds a reference value; d) decrementing a pumpspeed when the amplitude of the second harmonic term of the Fourierseries exceeds the reference value; e) repeating steps a-d until theamplitude of the second harmonic term falls below the reference value.201. A pump system for pumping blood, the system comprising: a bloodpump comprising: a) an inlet, b) an outlet, c) an impeller locatedfluidly between the inlet and outlet, and d) an impeller drive; and acontrol system comprising: a) an outflow pressure sensor located to reada pressure of blood flowing downstream of the impeller, b) an inflowpressure sensor located to read a pressure of blood flowing upstream ofthe impeller, c) a processor in communication with the outflow pressuresensor and inflow pressure, and d) a memory with at least one storedtarget pressure difference, wherein the processor reads both a pressurereading from the outflow pressure sensor and a pressure reading from theinflow pressure sensor and calculates a pressure difference, and whereinthe processor causes the pump impeller drive to operate so thecalculated pressure difference is within an acceptable range of the atleast one stored target pressure difference.
 202. A centrifugal bloodpump system comprising: a blood pump comprising: a pump housing havingan inlet cap, a top impeller casing, and a bottom impeller casing wherethe inlet cap engaged with the top impeller casing defines an inlet andwhere the top impeller casing engaged with the bottom impeller casingdefines an outlet, the inlet cap further defining a top recess toreceive a top bearing pin, and the bottom impeller casing furtherdefining a bottom recess to receive a bottom bearing pin; the topbearing pin and the bottom bearing pin each having a hemisphericalconcave bearing surface on a first end and a generally planar surface ona second opposite end, each hemispherical concave bearing surface havinga minimum diameter between 1 mm and 3 mm with a radius between 0.2 mmand 0.6 mm, the top bearing pin and the bottom bearing pin each having alength between about 6 mm and 8 mm and a circumferentially groove toprovide a mechanical interlock with the top recess and the bottomrecess, respectively; an impeller suspended within the pump housing, theimpeller defining: a pivot lumen to receive an impeller pivot, the pivotlumen extending along the central axis of the impeller between the toprecess and the bottom recess and the impeller pivot having a convexfirst end to engage the concave bearing surface of the top bearing pinand a convex second end to engage the concave bearing surface of thebottom bearing pin, the impeller pivot having an overall length between8 mm and 12 mm and a pivot diameter between 1 mm and 3 mm, where theconvex first end has a radius between 0.3 mm and 0.7 mm and extends awayfrom the impeller pivot a length between 0.4 mm and 0.6 mm, and wherethe convex second end has a radius between 0.2 mm and 0.4 mm and extendsaway from the impeller pivot a length between 0.4 mm and 0.6 mm; aplurality of arcuate blades on the top surface of the impeller andextending radially from a center of the impeller to an outer edge of theimpeller, the arcuate blades to force blood received at the inletthrough the pump housing and to the outlet; and at least one washoutlumen extending parallel to a central axis of the impeller from the topsurface through the impeller to a bottom surface; at least one magnetmechanically engaged to the impeller; an electric motor to magneticallyengage the at least one magnet, wherein the electric motor rotates theat least one magnet and the impeller; and a catheter system comprising:an inflow catheter having an distal inflow end and an proximal inflowend, the proximal inflow end fluidly connected to the inlet; and, anoutflow catheter having a distal outflow end and a proximal outflow end,the proximal outflow end fluidly connected to the outlet.
 203. Asensorless method of estimating and controlling a wall shear stress in ablood vessel comprising: a) estimating a blood viscosity; b) measuringat least one motor state variable of a blood pump system selected from avoltage, a current, or a pump speed; c) estimating a blood flow rate anda pressure head in the blood pump system; d) determining a vascularresistance of the blood vessel from the estimated blood flow rate andthe estimated pressure head; e) estimating a radius of the blood vessel;f) determining the wall shear stress from the estimated blood viscosity,the estimated blood flow rate, and the radius of the blood vessel; g)comparing the determined wall shear stress with a predeterminedreference value; h) adjusting the pump speed when the determined wallshear stress does not approximate the predetermined reference value; andi) repeating steps a-h until the determined wall shear stressapproximates the predetermined reference value.